Radiological imaging apparatus

ABSTRACT

A radiological imaging apparatus of the present invention comprises an image pickup device and a medical examinee holding device that is provided with a bed. The image pickup device includes a large number of radiation detectors and radiation detector support plates. A large number of radiation detectors are mounted around the circumference of a through-hole and arranged in the axial direction of the through-hole. The radiation detectors are arranged in three layers formed radially with respect to the center of the through-hole and mounted on the lateral surfaces of the radiation detector support plates. Since the radiation detectors are not only arranged in the axial direction and circumferential direction of the through-hole but also arrayed in the radial direction, it is possible to obtain accurate information about a γ-ray arrival position in the radial direction of the through-hole (the positional information about a radiation detector from which a γ-ray image pickup signal is output). The use of accurate information about γ-ray arrival increases the tomogram accuracy. As a result, the present invention enhances the tomogram accuracy, that is, the PET examination accuracy.

BACKGROUND OF THE INVENTION

[0001] The present invention relates to a radiological imagingapparatus, and more particularly to a radiological imaging apparatusideally applicable to X-ray computed tomography, positron emissioncomputed tomography (hereinafter referred to as “PET”), single-photonemission computed tomography (hereinafter referred to as “SPECT”),digital X-ray examination flat panel detector, and similar equipment.

[0002] Radiological imaging is a non-invasive imaging technology toexamine physical functions and conformation of a medical examinee as asubject. Typical radiological imaging devices are X-ray computedtomography, digital X-ray examination, PET, and SPECT devices.

[0003] PET is a method for administering a radiopharmaceutical(hereinafter referred to as a “PET pharmaceutical”) containingpositron-emitting nuclides (¹⁵O, ¹³N, ¹¹C, ¹⁸F, etc.), which areradionuclides, to a medical examinee, and examining locations in theexaminee's body where the PET pharmaceutical is heavily consumed. Morespecifically, the PET method is used to detect γ-rays that are emittedfrom the medical examinee's body due to the administered PETpharmaceutical. A positron emitted from the radionuclides contained inthe PET pharmaceutical couples with an electron of a neighboring cell(cancer cell) to disappear, emitting a pair of γ-rays (paired γ-rays)having an energy of 511 keV. These γ-rays are emitted in directionsopposite to each other (180°±0.6°). Detecting this pair of γ-rays usingradiation detectors makes it possible to locate the two radiationdetectors between which positrons are emitted. Detecting many of theseγ-ray pairs makes it possible to identify the locations where the PETpharmaceutical is heavily consumed. For example, when a PETpharmaceutical produced by combining positron-emitting nuclides withglucose is used, it is possible to locate carcinomatous lesions havinghyperactive glucose metabolism. The data obtained is converted toindividual voxel data by the filtered back projection method, which isdescribed on pages 228 and 229 of IEEE Transaction on Nuclear Science,Vol. 21. The half-life period of positron-emitting nuclides (¹⁵O, ¹³N,¹¹C, ¹⁸F, etc.) used for PET examination ranges from 2 to 110 minutes.

[0004] In PET examination, γ-rays generated upon positron annihilationattenuate within the human body so that transmission data is imaged tocompensate for γ-ray attenuation within the human body. Transmissiondata imaging is a method of measuring the γ-ray attenuation within themedical examinee's body by, for instance, allowing γ-rays to enter theexaminee's body using cesium as a radiation source and measuring theradiation intensity prevailing after penetration through the examinee'sbody. The PET image accuracy can be enhanced by estimating the γ-rayattenuation within the examinee's body from the measured γ-rayattenuation rate and correcting the data derived from PET examination.

[0005] A method for increasing the PET examination accuracy is describedon page 15 of Medical Imaging Technology, Vol. 18-1. This method is usedto insert a reflection plate into a crystal, acquire the informationabout depth with a DOI (Depth-Of-Interaction) detector, and reconstructthe image according to the acquired information to improve the imagequality. For the use of this method, it is necessary to use a radiationdetector that is capable of acquiring the information about radiationdetector's position in the direction of the depth.

[0006] However, the use of a DOI detector involves image deterioration,which is caused by a decrease in the amount of signal transmissionsubstance. When, for instance, a 5 mm square BGO scintillator is used,approximately 200 photons are generated to function as a signaltransmission substance when there is a 511 keV incident γ-ray . However,when photons are partly reflected by a reflection plate as in the use ofthe DOI detector noted above, the amount of signal transmissionsubstance decreases. When the quantity of signal transmission substancereaching a photomultiplier tube is N and the incident γ-ray energy is E,the energy spectrum spread σ can be expressed by equation (1).

σ=E/{square root}N  (1)

[0007] Therefore, when the value N becomes smaller, the value σincreases to spread the energy spectrum. When the energy spectrum isspread, the correlation between the incident γ-ray energy and the signalgenerated by a DOI detector is impaired. As a result, this makes itdifficult to accurately measure the incident γ-ray energy.

[0008] If incident γ-ray energy measurements cannot be accurately made,it is difficult to remove scattered radiation contained in incidentγ-rays. In PET, the signal output from a radiation detector is passedthrough an energy filter for scattered radiation removal so as to detectonly γ-rays that have a specific energy level or higher. However, if theenergy spectrum is spread and, for example, the radiation detectorsignal output generated by 511 keV γ-rays cannot be differentiated fromthe radiation detector signal output generated by 300 keV γ-rays, it isnecessary to use an energy filter rated at 300 keV or lower. In thisinstance, 300 keV or higher scattered radiation is also measured so thatthe amount of noise increases. This can cause PET image deterioration.

[0009] SPECT is a method for administering a radiopharmaceutical(hereinafter referred to as a “SPECT pharmaceutical”) containingsingle-photon-emitting nuclides (⁹⁹Tc, ⁶⁷Ga, ²⁰¹Tl, etc.), which areradionuclides, and glucose or other substance that gathers aroundspecific tumors or molecules, to a medical examinee, and detecting aγ-ray emission from radionuclides with a radiation detector. The energyof γ-ray emission from single-photon-emitting nuclides, which arefrequently used for SPECT examination, is approximately several hundredkeV. In SPECT, a single γ-ray is emitted so that the angle of γ-rayincidence upon a radiation detector cannot be determined. Therefore, acollimator is used to obtain angular information by detecting only theγ-radiation incident at a specific angle. The SPECT is an examinationmethod for detecting γ-rays generated within a medical examinee's bodydue to the SPECT pharmaceutical for the purpose of identifying thelocations where the SPECT pharmaceutical is heavily consumed. The dataobtained is converted to individual voxel data by the filtered backprojection or like method as is the case with PET. It should be notedthat transmission images may also be generated in SPECT. The half-lifeperiod of ⁹⁹Tc, ⁶⁷Ga, and ²⁰¹Tl, which are used for SPECT, is longerthan that of PET radionuclides and from 6 hours to 3 days.

[0010] X-ray CT (computed tomography) is a method for exposing a medicalexaminee to radiation emitted from a radiation source and imaging theconformation within the examinee's body in accordance with radiationtransmittance in the examinee's body. The intensity of X-rays passingthrough the examinee's body, which is measured with a radiationdetector, is used to determine the coefficient of linear attenuationwithin the examinee's body between the X-ray source and radiationdetector. The determined linear attenuation coefficient is used todetermine the linear attenuation coefficient of each voxel by theaforementioned filtered back projection method. The resulting value isthen converted to a CT value.

[0011] A flat panel detector is a flat radiation detector for use indigital X-ray examination, which is a digital version of conventionalX-ray examination. Being equipped with such a flat radiation detectorinstead of a conventional X-ray film, a flat panel detector imagingdevice detects X-rays passing through a medical examinee's body, handlesthe information about attenuation within the examinee's body as digitalinformation, and displays the digital information on a monitor. The flatpanel detector imaging device does not require the use of X-ray film orother media and displays an image immediately after image exposure.

[0012] For maintenance of examination accuracy, all these radiologicalimaging apparatuses require their radiation detectors to be subjected todetection efficiency calibration at periodic intervals of, for instance,three months. The radiation detector's detection efficiency deteriorateswith time. However, the deterioration characteristic varies from oneradiation detector to another. It is therefore necessary to determinethe detection efficiency of each radiation detector on a periodic basis.In PET or SPECT examination in which the number of photons incident oneach radiation detector is measured, correct measurement cannot be madeif the detection efficiency varies from one radiation detector toanother. Therefore, the detection efficiency of each radiation detectoris determined beforehand, and the value of each radiation detector ismultiplied by the reciprocal of the determined detection efficiencyvalue in order to compensate for image deterioration resulting from thedetection efficiency variation of radiation detectors. In X-ray CT orflat panel detector examination, on the other hand, the X-ray intensityis detected by radiation detectors; however, intensity measurements needto be corrected if the detection efficiency varies.

[0013] As explained above, the use of radiological imaging apparatusesentails an enormous amount of time and labor because they require theirradiation detectors to be checked for detection efficiency variation inorder to maintain examination accuracy.

SUMMARY OF THE INVENTION

[0014] It is an object of the present invention to provide aradiological imaging apparatus that determines the locations reached byradiation with increased precision and enhances the accuracy of imagesto be generated.

[0015] The present invention to attain the above-described object ischaracterized by comprising an image pickup device, which comprises aplurality of radiation detectors for detecting radiation from a subject,wherein the radiation passing through a first radiation detector is tobe detected by a second radiation detector. Since the second radiationdetector is provided to detect the radiation passing through the firstradiation detector, the locations reached by radiation (locations atwhich radiation is detected) can be confirmed, with increased precision,in the direction of the depth from the first radiation detector opposingthe subject. As a result, a highly accurate image depicting the interiorof the subject's body can be obtained.

[0016] Preferably, the present invention comprises a plurality ofradiation detectors that enable an image pickup device to detectradiation from a subject, wherein the radiation detectors are formed inthe image pickup device and positioned around the circumference of thethrough-hole, into which a bed is to be inserted, and at differentradial locations.

[0017] Preferably, the present invention also comprises a plurality ofradiation detectors that enable an image pickup device to detectradiation from a subject, wherein the radiation detectors are mounted onradiation detector support members that are positioned around thecircumference of the through-hole, into which a bed is to be inserted,and at different radial locations.

[0018] In addition, the present invention attaining the above-describedobject is characterized by comprising a plurality of radiation detectorsfor γ-ray detection, wherein a radiation detector detecting unscatteredinternal γ-rays is located within a preselected period of time and inaccordance with the detection signals output from at least threeradiation detectors and the position information about these radiationdetectors.

[0019] The present invention makes it possible to find the sequence ofunscattered γ-ray attenuation (scatter sequence) in accordance withthree or more detection signals output within a preselected period oftime and the positional information about three or more radiationdetectors that generated the detection signals, and determine theposition and direction of γ-ray initial incidence. In marked contrast todetermining the γ-ray initial incidence position in a random manner, thepresent invention is capable of locating unscattered γ radiation withhigh efficiency and generating highly accurate tomograms.

[0020] In addition, the present invention attaining the above-describedobject is characterized in that it comprises a plurality of radiationdetectors for γ-ray detection, and that when detection signals areoutput from at least three of such radiation detectors within apreselected period of time, the attenuation sequence, initial incidenceposition, and initial incidence direction of one of paired γ-rays aredetermined in accordance with the positional information about at leasttwo of such radiation detectors, the energy detection values of at leasttwo of such radiation detectors, and the positional information about aradiation detector detecting the remaining γ radiation of paired γ-rays.

[0021] The present invention determines the attenuation sequence(scatter sequence) of one of paired γ-rays in accordance with thepositional information about the remaining paired γ-ray , and determinesthe position and direction of γ-ray initial incidence on a radiationdetector. More specifically, the positional information about eachradiation detector detecting a first one of paired γ-rays and thepositional information about a radiation detector detecting theremaining paired γ-ray are used to estimate two or more possibleattenuation sequences of the first one of the paired γ-rays. Theestimated attenuation sequences are examined to choose the one thatexhibits the proper correlation between the scatter angle and energydetection value of the first one of the paired γ-rays. The γ-rayattenuation sequence is determined in this manner. As a result, theposition of initial γ-ray incidence on a radiation detector (theposition of a radiation detector related to the first γ-ray attenuation)is determined. Consequently, it is possible to conclude that a γ-raygeneration source (diseased area) exists on a straight line (directionof initial incidence) joining the located radiation detector and theradiation detector detecting the remaining paired γ-ray . In markedcontrast to determining the γ-ray initial incidence position in a randommanner, the present invention is therefore capable of locatingunscattered γ radiation with high efficiency and generating highlyaccurate PET images.

[0022] In addition, the present invention attaining the above-describedobject is characterized in that it comprises a plurality of radiationdetectors for γ-ray detection and collimators mounted in front of theradiation detectors to permit γ-ray passage, and that when detectionsignals are output from at least three of such radiation detectorswithin a preselected period of time, the attenuation sequence, initialincidence position, and initial incidence direction of γ radiation aredetermined in accordance with the positional information about at leastthree of such radiation detectors and the energy detection values of atleast three of such radiation detectors.

[0023] When the detection signals of three or more radiation detectorsare simultaneously counted (output within the specified period of time),the present invention determines the γ-ray attenuation sequence (scattersequence) in accordance with the positional information about the threeor more detection signals and the energy detection values from the threeor more radiation detectors, and determines the position and directionof γ-ray incidence on a radiation detector. More specifically, theabove-mentioned positional information is first used to estimate two ormore possible sequences of γ-ray attenuation, and the estimatedsequences are checked to choose the one that exhibits the propercorrelation with the above-mentioned energy detection values. Inaccordance with the determined γ-ray initial incidence position and theabove energy detection values, the direction of γ-ray initial incidencecan be determined. In marked contrast to determining the γ-ray initialincidence position in a random manner, the present invention istherefore capable of locating unscattered γ radiation with highefficiency and generating highly accurate PET images.

BRIEF DESCRIPTION OF THE DRAWINGS

[0024]FIG. 1 is a configuration diagram showing a radiological imagingapparatus according to a preferred embodiment of the present invention;

[0025]FIG. 2 is a cross sectional view taken along line II-II of FIG. 1;

[0026]FIG. 3 is a perspective view illustrating the structure of aradiation detector support shown in FIG. 1;

[0027]FIG. 4A is a longitudinal sectional view of a calibrated radiationsource shown in FIG. 1;

[0028]FIG. 4B is a cross sectional view taken along line IV-IV of FIG.4A;

[0029]FIG. 5 is a flow chart illustrating a tomogram generation processthat is performed by a computer shown in FIG. 1;

[0030]FIG. 6 is a flow chart that details the process performed in step42 shown in FIG. 5;

[0031]FIG. 7 presents diagrams that indicate how γ-rays are detected inan embodiment shown in FIG. 1;

[0032]FIG. 8 is a configuration diagram showing a radiological imagingapparatus (SPECT examination apparatus) according to another embodimentof the present invention;

[0033]FIG. 9A is a longitudinal sectional view of a calibrated radiationsource shown in FIG. 8;

[0034]FIG. 9B is a cross sectional view taken along line IX-IX of FIG.9A;

[0035]FIG. 10 is a configuration diagram showing a radiological imagingapparatus according to another embodiment of the present invention;

[0036]FIG. 11 is a diagram illustrating a radiation detector-to-signalprocessor connection according to the embodiment shown in FIG. 10;

[0037]FIG. 12 is a configuration diagram of a signal discriminator shownin FIG. 11;

[0038]FIG. 13 is a flow chart illustrating a tomogram generation processthat is performed by a computer shown in FIG. 10;

[0039]FIG. 14 is a configuration diagram showing a radiological imagingapparatus according to another embodiment of the present invention;

[0040]FIG. 15 is a diagram illustrating a typical radiation detectorarrangement for a flat panel display shown in FIG. 14;

[0041]FIG. 16 is a configuration diagram showing a radiological imagingapparatus according to another embodiment of the present invention;

[0042]FIG. 17 is a configuration diagram showing a radiological imagingapparatus according to another embodiment of the present invention;

[0043]FIG. 18 is a diagram illustrating a radiation detector-to-signalprocessor connection according to the embodiment shown in FIG. 17;

[0044]FIG. 19 is a cross sectional view taken along line X-X of FIG. 17;

[0045]FIG. 20 is a flow chart illustrating a tomogram generation processthat is performed by a computer shown in FIG. 17;

[0046]FIG. 21 shows characteristic curves that indicate the γ-rayenergy-scatter angle relationship prevailing before and afterscattering;

[0047]FIG. 22 is a flow chart that illustrates how a coincidence countershown in FIG. 18 determines the position and direction of γ-ray initialincidence;

[0048]FIG. 23 is a configuration diagram showing a radiological imagingapparatus according to another embodiment of the present invention;

[0049]FIG. 24 is a cross sectional view taken along line Y-Y of FIG. 23;and

[0050]FIG. 25 shows a typical signal input/output of a coincidencecounter according to the embodiment shown in FIG. 23.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0051] Embodiment 1

[0052] A radiological imaging apparatus according to a preferredembodiment of the present invention will be described below withreference to FIGS. 1 and 2. A radiological imaging apparatus 1 ofEmbodiment 1 is used for PET examination. This apparatus comprises animage pickup device 2, a signal processor 7, a tomogram generator 10, amedical examinee-holding device 14, a calibrated radiation sourcecircumferential transfer unit 37, and a drive controller 35.

[0053] The image pick device 2 has a casing 3, a large number ofradiation detectors 4, and a large number of radiation detector supportplates 5. The casing 3 has an opening (through-hole) 6 into which amedical examinee or a subject is to be inserted. A large number of theradiation detectors (e.g., 10,000 radiation detectors in total) 4 arepositioned around the circumference of the through-hole 6 and arrangedin the axial direction of the through-hole 6. As shown in FIG. 2, theinnermost radiation detectors 4 are circularly disposed around thecircumference of the through-hole 6. The other radiation detectors 4 arearranged radially from the center of the through-hole 6 with theabove-mentioned innermost radiation detectors 4 regarded as the startingpoints. The radiation detectors 4 are also disposed at their respectivedifferent positions in the radial directions of the through-hole 6. Thatis, Embodiment 1 is configured so that multiple sets of three radiationdetectors 3 (e.g., the radiation detectors 4 a, 4 b, and 4 c shown inFIG. 2) are linearly positioned to form three layers in the radialdirection of the through-hole 6. Each layer of radiation detectors 4 ispositioned in a circular form (e.g., concentrically).

[0054] As shown in FIG. 3, the radiation detectors 4 are mounted on alateral surface of the radiation detector support plate 5. Morespecifically, the radiation detectors 4 are radially mounted on alateral surface of the radiation detector support plate 5, which isshaped like a half ring. A plurality of radiation detector supportplates 5, on which the radiation detectors are mounted, are mounted onthe lower inner surface of the through-hole 6 and arranged in the axialdirection of the through-hole 6. These radiation detector support plates5 are fastened to the casing 3. Although not shown in FIG. 3, aplurality of radiation detector support plates on which the radiationdetectors 4 are mounted, are also mounted on the upper inner surface ofthe through-hole 6, arranged in the axial direction of the through-hole6, and fastened to the casing 3. One radiation detector support plate 5mounted on the lower inner surface of the through-hole 6 and oneradiation detector support plate 5 mounted on the upper inner surface ofthe through-hole 6 are positioned so as to form a ring in the sameplane. The radiation detector support plates 5 may also be shaped in acircular form.

[0055] The signal processor 7 comprises γ-ray discriminators 8 and acoincidence counter 9 provided for each of the radiation detectors 4.The γ-ray discriminators 8 are connected to their respective radiationdetectors 4 via a wiring 13. The number of installed γ-raydiscriminators 8 is equal to that of installed radiation detectors 4.The coincidence counter 9 is connected to all the γ-ray discriminators8. The tomogram generator 10 comprises a computer 11, a storage device12, and a display device 130. The computer 11 is connected to thecoincidence counter 9. The storage device 12 is connected to thecomputer 11. The display device is also connected to the computer 11.The medical examinee-holding device 14 is provided with a support 15 anda bed 16, which is mounted on the top of the support 15 so as to bemoved in the longitudinal direction. The image pickup device 2 isdisposed in a direction perpendicular to the longitudinal direction ofthe bed 16.

[0056] Typical examples of a radiation detector include a semiconductorradiation detector and a scintillator. The scintillator is not suitablefor a multilayer arrangement (e.g., aforementioned three layers) becausea photomultiplier or like device needs to be mounted on the rear of acrystal (BGO, NaI, etc.), which serves as a radiation detector. On theother hand, the semiconductor radiation detector is suitable for amultilayer arrangement because it does not require the use of aphotomultiplier or like device. In Embodiment 1, semiconductor radiationdetectors are used as the radiation detectors 4, and their detectionunit, which is a 5 mm cube, is made of cadmium telluride (CdTe). Thedetection unit may also be made of gallium arsenide (GaAs) or cadmiumzinc telluride (CZT).

[0057] The calibrated radiation source circumferential transfer unit 37includes a guide rail 28 and a calibrated radiation source device 29.The guide rail 28 is circular, mounted on a lateral surface of themedical examinee holding device 14 on the casing, and arranged aroundthe circumference of the through-hole 6. The calibrated radiation sourcedevice 29 has a calibrated radiation source drive 30 and a calibratedradiation source 31. The calibrated radiation source drive 30 is movablymounted on the guide rail 28. Although not shown in the figure, thecalibrated radiation source drive 30 includes a pinion that engages witha rack on the guide rail 28, and a motor that rotates the pinion via aspeed reduction mechanism. The calibrated radiation source 31 is mountedon the casing (not shown) for the calibrated radiation source drive 30and attached to the distal end of an arm 38 that is horizontallytelescopic. As shown in FIGS. 4A and 4B, the calibrated radiation source31 houses a γ-ray source 33 within a γ-ray shield 32 having aunidirectional opening. Except the above-mentioned opening, the externalsurface of the γ-ray shield 32 is covered by a casing (not shown)serving as an enclosure. The calibrated radiation source 31 includes amovable shutter 34, which is capable of covering the opening in theγ-ray shield 32. A Ga—Ge radiation source for 511 keV γ-ray emissions isused as the γ-ray source 33. A Cs radiation source for 662 keV γ-rayemissions may be used instead of the Ga—Ge radiation source. Thecalibrated radiation source 31 is a radiation source for use duringtransmission data imaging. A collimator 39 that is positioned in frontof the opening in the γ-ray shield 32 is mounted on the γ-ray shield 32so as not to obstruct the open/close operation of the shutter 34.

[0058] First of all, transmission data imaging by a radiological imageapparatus 1A will be described. Transmission data imaging is a techniquefor measuring the γ-ray transmittance with a medical examinee's bodywith a calibrated radiation source. The time required for measurement isabout 1 or 2 minutes. After γ-rays emitted from the calibrated radiationsource pass through a medical examinee, they are measured by radiationdetectors 4. The rate of γ-ray attenuation within the medical examinee'sbody is determined in accordance with the radiation intensity of thecalibrated radiation source and the measured γ radiation. The determinedγ-ray attenuation rate is used to compensate for an in-vivo scatter(phenomenon in which γ-rays generated within a medical examinee's bodydue to a radiopharmaceutical are scattered and attenuated) during PETexamination.

[0059] The details of transmission data imaging will be described below.The medical examinee 17 laid on the bed 16 is inserted into thethrough-hole 6. When transmission data imaging starts, the radiationsource controller 69 opens the shutter 34. γ-rays emitted from the γ-raysource 33 pass through the opening in the γ-ray shield 32 and collimator39, and then fall on the medical examinee 17. The directivity of γ-raysemitted from the γ-ray source 33 is increased by the collimator 39 sothat the direction of γ-ray travel is determined. At the beginning oftransmission data imaging, the drive controller 35 outputs a drive startsignal to rotate the motor of the calibrated radiation source drive 30.When the motor rotates, the calibrated radiation source drive 30 moveson a guide rail 28 to circulate around the medical examinee 17. Withinthe through-hole 6, the calibrated radiation source 31 moves around themedical examinee 17. Therefore, highly directional γ-rays emitted fromthe calibrated radiation source 31 are incident on the medical examinee17 from all circumferential positions. The bed 16 moves toward theopposite end of the through-hole 6. After γ-rays pass through themedical examinee 17, they are measured by the radiation detectors 4.Since highly directional γ-rays are emitted, unscattered γ-rays aremeasured by the radiation detectors 4. These rays have the same 511 keVenergy as when they are emitted from the γ-ray source 33.

[0060] The radiation detectors 4 measure the γ-rays passing through themedical examinee 17 and output a γ-ray detection signal. In response tothis γ-ray detection signal, the γ-ray signal discriminator 8 generatesa pulse signal as is the case with the γ-ray detection signal detectedduring PET examination described later. The coincidence counter 9measures the pulse signal and outputs its count and the two points ofpaired γ-ray detection (the positions of a pair of radiation detectors 4that are mounted about 180° apart from each other with respect to theaxial center of the through-hole 6). The computer 11 stores the countand the positional information about the two detection points in thestorage device 12. At the end of transmission data imaging, the drivecontroller 35 outputs a drive end signal to stop the motor of thecalibrated radiation source drive 30. At this time, the radiation sourcecontroller 69 closes the shutter 34 of the calibrated radiation source31 so as to prevent γ-rays from being emitted outside.

[0061] Three radiation detectors that are linearly arranged in thedirection of the radius of the through-hole 6 to form three layers(e.g., radiation detectors 4 a, 4 b, and 4 c shown in FIG. 1) arehandled as a radiation detector group. Embodiment 1 provides a pluralityof radiation detector groups. When the energy of emitted γ-rays isuniform, the γ-ray detection efficiency is determined by a theoreticalformula. Since the radiation detectors 4 are semiconductor radiationdetectors having a detection unit made of 5 mm thick CdTe, the detectionefficiency of 511 keV γ-rays is about 20%. In a single radiationdetector group, therefore, incident γ-radiation is attenuated by about20% in the first layer radiation detector 4, and the 80% γ-radiationpassing through the first layer radiation detector 4 is attenuated byabout 20% in the second layer radiation detector 4, that is, about 16%γ-radiation attenuation occurs in the second layer radiation detector 4.In the third layer radiation detector 4, the 64% γ-radiation passingthrough the second layer radiation detector 4 is attenuated by about20%, that is, about 12.8% γ-radiation attenuation occurs. γ-raydetection signals reflecting such attenuations are output from the firstand second layer radiation detectors 4. These γ-ray detection signalsare fed to the γ-ray discriminators 8 of the associated signalprocessors 7, subjected to a scattered γ-ray removal process, andconverted to pulse signals. The coincidence counters 9 of the signalprocessors 7 measure the pulse signals. When the γ-ray detection signalsfed from the layered radiation detectors 4 are independently measuredand the measurement result significantly differs (by, for instance, morethan ±5%) from the theoretical detection efficiency proportion (approx.20:16:12.8) of the first-layer to third-layer radiation detectors 4, itmeans that the detection efficiency of one or more of the radiationdetectors 4 is decreased due to radiation detector deterioration. If,for instance, one radiation detector 4 is deteriorated and the other tworadiation detectors 4 are operating normally, the measured detectionefficiency proportion of the affected radiation detector group greatlydiffers from the above-mentioned theoretical value. It is thereforepossible to locate the radiation detector 4 that is deteriorated.Further, the percentage of detection efficiency decrease caused bydeterioration can be calculated from the detection efficiencies,detection efficiencies determined from the above-mentioned proportion,and measured detection efficiencies of the two normal radiationdetectors 4. When, for instance, the measured detection efficiencyproportion determined from the measurements of three radiation detectorsin one radiation detector group is 20:4:12.8, the measured detectionefficiency of the second layer radiation detector 4 is 12 points lower(75% lower) than the theoretical detection efficiency. It means that thesecond layer radiation detector is faulty.

[0062] The concept of fault detection will be described below. Whenγ-rays are emitted from the γ-ray source 33 at a certain time, they areincident on three radiation detectors 4 in one radiation detector group(e.g., radiation detectors 4 a, 4 b, and 4 c shown in FIG. 2) but not onthe three radiation detectors 4 in another radiation detector group(e.g., a radiation detector group adjacent to the first one). Thedetection efficiency proportion of the radiation detectors 4 in a singleradiation detector group is determined from the data indicating theprevious deterioration of the radiation detectors 4 while consideringthe γ-ray transmission distance and γ-ray transmission sequence.Further, the theoretical detection efficiency proportion of theradiation detectors 4 in the radiation detector group is determined byperforming a simulation or theoretical calculations. The measureddetection efficiency proportion determined according to the γ-raydetection signals generated from the radiation detectors 4 in theradiation detector group is compared with the above-mentionedtheoretical detection efficiency proportion to check whether or not theradiation detectors 4 in the radiation detector group are deteriorated.All the radiation detector groups are subjected to the comparisonbetween the measured detection efficiency proportion and theabove-mentioned theoretical detection efficiency proportion. If all theradiation detectors 4 mounted on the image pickup device 2 are of thesame type, the theoretical detection efficiency proportion can bedetermined by performing calculations on only one representativeradiation detector group. Further, the measured detection efficiencyproportion determined from the individual γ-ray detection signals iscompared with the detection efficiency proportion determined from thedata indicating the previous deterioration of the radiation detectors 4to check the progress of deterioration of the radiation detectors 4 inthe radiation detector group. If the radiation detectors 4 aredeteriorated, the storage device 12 stores the information about thedegree of deterioration, and the user is notified of deterioration andfault. When this process is repeated for all radiation detector groups,it is possible to grasp the degree of detection efficiency deteriorationof the individual radiation detectors mounted on the image pickup device2 and remove faulty radiation detectors. A specific process based on theabove-described concept of fault detection will be described later withreference to FIGS. 5 and 6.

[0063] Next, a PET examination process performed with the radiologicalimaging apparatus 1 will be described. A PET pharmaceutical is injectedinto or otherwise administered to the medical examinee 17 or the subjectin advance. The medical examinee 17 stands by for a predetermined periodof time so that the PET pharmaceutical is diffused within its body andgathered at a diseased area (e.g., carcinomatous lesion) to permitimaging. An appropriate PET pharmaceutical is selected in accordancewith a lesion to be checked for. When the predetermined period of timeelapses, the medical examinee 17 is laid on the bed 16 and subjected toPET examination with the image pickup device 2. When PET examinationstarts, the bed 16 moves toward the image pickup device 2 and goes intothe through-hole 6, carrying the medical examinee 17 into thethrough-hole 6. When 511 keV γ-rays (in the case where the PETpharmaceutical contains ¹⁸F) are emitted from a lesion in the body ofthe medical examinee 17, they are incident on the radiation detectors 4.The radiation detectors 4 detect γ-rays that are emitted from the lesiondue to PET pharmaceutical administration, and generate γ-ray detectionsignals. The γ-ray detection signals are delivered to the associatedγ-ray discriminators 8 via the associated wiring 13. Each γ-raydiscriminator includes a waveform shaper (not shown). The waveformshaper converts an input γ-ray detection signal to a γ-ray detectionhaving a time Gaussian distribution waveform. The energy of γ-raysgenerated upon annihilation (at the lesion) of positrons emitted fromthe PET pharmaceutical is 511 keV. However, if the γ-rays are scatteredwithin the body of the medical examinee, the energy is lower than 511keV. To remove scattered γ-rays, the γ-ray discriminator 8 includes afilter (not shown) for selecting an energy setting of, for instance, 400keV, which is lower than 511 keV, and passing γ-ray detection signalshaving a energy value greater than the energy setting. This filterreceives the γ-ray detection signals that are output from the waveformshaper. An energy setting of 400 keV is selected here as an example inconsideration of the variation of γ-ray detection signals generated upon511 keV γ-ray incidence on the radiation detectors 4. In response to aγ-ray detection signal passing through the filter, the γ-raydiscriminator 8 generates a pulse signal having a predetermined energylevel.

[0064] The coincidence counter 9 receives pulse signals generated by allthe γ-ray discriminators 8, and determines the count data for the γ-raydetection signals output from the radiation detectors 4. Further, thecoincidence counter 9 uses the pulse signals for the aforementionedpaired γ-rays to determine the positional information about two pointsof paired γ-ray detection. The positional information about thesedetection points is transmitted to the computer 11 and stored into thestorage device 12 by the computer 11. The count data for theabove-mentioned γ-ray detection signals are also stored into the storagedevice 12 by the computer 11.

[0065] The computer 11, using the count data and other relevant data,performs a processing procedure indicated in FIGS. 5 and 6 toreconstruct a tomogram of the medical examinee 17. The processingprocedure will be detailed below. The count data derived from PETexamination, the positional information about the associated detectionpoints, and the count data derived from transmission imaging are readfrom the storage device 12 and input (step 40). The theoreticaldetection efficiency proportion of the radiation detectors in aradiation detector group is calculated (step 41). This theoreticalproportion is determined by performing theoretical calculations on thetransmission distance of γ-rays emitted from the medical examinee 17(this distance varies with the radionuclides contained in the PETpharmaceutical) and the sequence of γ-ray transmission. When a PETpharmaceutical containing ¹⁸F is administered to the medical examinee17, the theoretical detection efficiency proportion of three radiationdetectors 4 in the radiation detector group is about 20:16:12.8. UnlikeEmbodiment 1 in which the theoretical detection efficiency proportion iscalculated each time, the theoretical detection efficiency proportion ofall the radiation detectors in the radiation detector group may becalculated in advance for various PET pharmaceuticals containingdifferent radionuclides and stored in the storage device 12.

[0066] Next, deteriorated radiation detectors 4 are checked for (step42). The process in step 42 is performed on a radiation detector groupbasis and will be described in detail with reference to FIG. 6. First, aradiation detector group is selected (step 50). The measured detectionefficiency proportion of radiation detectors in the selected radiationdetector group is calculated (step 51). More specifically, the countdata derived from γ-ray detection signals output from the radiationdetectors 4 in the selected radiation detector group are used tocalculate the measured detection efficiency proportion of the radiationdetectors. The difference between the measured detection efficiencyproportion and theoretical detection efficiency proportion is checked todetermine whether it is within a preselected range (±5% of thetheoretical proportion) (step 52). When the difference is within thepreselected range (when the answer to the question is “Yes”), theradiation detectors 4 in the selected radiation detector group areoperating normally without being deteriorated. However, if thedifference is outside the preselected range (when the answer to thequestion is “No”), the information about deteriorated radiationdetectors 4 in the radiation detector group (these detectors arereferred to as deteriorated radiation detectors) is stored in thestorage device 12 (step 53). When the difference is outside thepreselected range in this manner, it means that one or more radiationdetectors 4 in the radiation detector group are deteriorated.Deteriorated radiation detectors in the radiation detector group can belocated by comparing the measured detection efficiency proportion valuesof the radiation detectors with their theoretical ones as explainedearlier. Next, the deterioration information about a deterioratedradiation detector 4 is output to the display device 130 (step 54). Thedeterioration information about the deteriorated radiation detector 4 isthe detection efficiency proportion information that is obtained usingthe data indicating the degree of previous deterioration of theradiation detector 4. In accordance with the deterioration informationabout the deteriorated radiation detector 4, which is shown on thedisplay device 130, the operator is able to grasp the degree ofdeterioration of the deteriorated radiation detector 4 on the basis ofthe deterioration information about the deteriorated detector 4displayed on the display device 130. If a deteriorated radiationdetector 4 is significantly deteriorated, it needs to be replaced by anew radiation detector. The detection efficiency of the deterioratedradiation detector 4 is corrected (step 55). If, for instance, themeasured proportion values of radiation detectors 4 a and 4 c in aradiation detector group coincide with the theoretical ones and themeasured proportion value of radiation detector 4 b is considerablysmaller than the theoretical one, the detection efficiency of radiationdetector 4 b is corrected to a detection efficiency that can beestimated from the measured detection efficiency proportion of radiationdetectors 4 a and 4 c and the theoretical detection efficiencyproportion of radiation detectors 4 a, 4 b, and 4 c. The count datadetermined according to the corrected detection efficiency is stored inthe storage device 12 as the count data about the radiation detector 4b.

[0067] When the answer to the question in step 52 is “Yes” or when theprocess in step 55 is terminated, it is checked whether or not “anyradiation detector groups remain to be selected” (step 56). When theanswer to the question in step 56 is “Yes”, the next radiation detectorgroup is selected in step 57. The processing from steps 51 onward isperformed until the answer to the question in step 56 changes to “No”.When the answer to the question in step 56 is “No”, a transmission imageis created (step 43). More specifically, the count data for γ-raydetection signals obtained at the time of transmission data imaging isused to calculate the γ-ray attenuation rate of each voxel in the bodyof the medical examinee 17. The γ-ray attenuation rate of each voxel isstored in the storage device 12.

[0068] Next, the in-vivo attenuation correction count for radiationdetectors is calculated (step 44). Since paired γ-rays is emitted duringthe PET examination, the in-vivo attenuation correction count iscalculated according to the sum of paired γ-ray move distances withinthe body. The count data derived from the PET examination, thepositional information about detection points, the γ-ray attenuationrate calculated in step 43 are used to reconstruct the tomogram of themedical examinee 17 by the tomogram reconstruction method describedlater in step 47. First of all, the γ-ray attenuation rate of eachvoxel, which is obtained in step 43, is used to determine the rate ofγ-ray attenuation between a pair of radiation detectors 4 (e.g.,radiation detectors 4 f and 4 g shown in FIG. 7B), which detects pairedγ-rays, according to the forward projection method. The reciprocal ofthe determined γ-ray attenuation rate is the attenuation correctioncount. In step 45, the attenuation correction count is used to providean in-vivo attenuation correction. The count data derived from the PETexamination is corrected by multiplying it by the attenuation correctioncount. Although the γ-rays generated at a lesion in the medical examinee17 are absorbed and attenuated during its transmission through the body,the accuracy of the count data derived from the PET examination can beincreased by correcting the count data with the above attenuationcorrection count.

[0069] In step 46, a γ-ray detection correction is also made inaccordance with the detection efficiency difference among radiationdetectors. Since paired γ-rays are emitted during the PET examination,the count data needs to be corrected using the detection efficiencies oftwo radiation detector groups at which respective paired γ-rays arrive.More specifically, the correction is made by multiplying the detectionefficiency correction counts of radiation detectors that have detectedγ-rays in the two radiation detector groups. This correction processwill be detailed below. The difference between the theoretical andmeasured detection efficiencies of each radiation detector 4, whichprevails during forward projection imaging, is determined in step 42.Let the theoretical detection efficiency value of the i-th radiationdetector 4 in radiation detector group j, which prevails during forwardprojection imaging, be Xfi_(ij), and the count data corrected in step 45be Xse_(ij). If the i-th detector is found to be faulty while the k-thdetector is normal, the corrected PET count data Xsi_(ij) for the i-thradiation detector is as expressed by equation (2). In order from theradiation detector 4 nearest the through-hole 6 to the radiationdetector 4 farthest from the through-hole 6, the value i is 1, 2, 3, andso son.

Xsi _(ij) =Xse _(kj) ×Xfi _(ij) /Xfi _(kj)  (2)

[0070] The corrected PET count data (count data corrected in accordancewith the detection efficiency difference among radiation detectors),which is calculated from equation (2), is stored in the storage device12.

[0071] The tomogram of the medical examinee, which contains a diseasedarea (e.g., carcinomatous lesion), is reconstructed (step 47). In step47, the tomogram is reconstructed using the corrected PET count dataXsi_(ij), which is derived from the correction made in step 46, and thepositional information about detection points. Tomogram reconstructionwill be detailed below. The computer 11 performs a tomogramreconstruction process with the above count data and detection pointpositional information while using the filtered back projection method.The computer 11 is a device for tomogram reconstruction. When thefiltered back projection method is used, the tomogram is reconstructedusing data that is sorted according to two parameters (distance t andangle θ), as described in the aforementioned document. The distance tand angle θ will be detailed with reference to FIG. 2. Suppose that thepaired γ-rays emitted from a lesion in the medical examinee 17 isdetected by radiation detectors 4 d and 4 e. A straight line 19 passesthrough the middle point of a line 18, which joins radiation detectors 4d and 4 e, and is at right angles to the line 18. The angle formedbetween a reference axis 20 (straight line oriented in any direction andpassing through the central point of a circle on which innermostradiation detectors are positioned, that is, the central point of thethrough-hole 6) and the line 19 is θ. The distance between the centralpoint 21 of the through-hole 6 and line 18 is t. The angle θ representsthe angle of rotation of the line 18 that joins radiation detectors 4 dand 4 e, which have detected paired γ-rays, from the reference axis 20.

[0072] In the radiological imaging apparatus 1, a plurality of radiationdetectors 4 is layered in the radial direction of the through-hole 6.Thanks to this layered arrangement, a new function described below canbe exercised. For example, suppose that two γ-rays 23 a, 23 b, which aregenerated at an γ-ray generation point 22 (lesion) in the body of themedical examinee 17 as shown in FIG. 7A, are incident on radiationdetectors 4 f and 4 g. Since the attenuation positions within thedetectors are unknown, a line joining the ends of a pair of radiationdetectors 4 f, 4 h, that is, a line 24 shown in FIG. 7B is regarded as adetection line when a conventional method is used. The radiologicalimaging apparatus 1, on the other hand, layers radiation detectors 4 inthe radial direction of the through-hole 6. Therefore, the γ-raydetection signal of radiation detector 4 g, which is located outward inthe radial direction, is obtained so that a line 25, which joinsradiation detectors 4 f and 4 g, can be used as a detection line. Inother words, the attenuation position in the direction of a detectordepth can be determined although it could not be determined through theuse of a conventional detector. As a result, the image accuracy isincreased because the detection line 25 accurately passes a location ofpaired γ-ray generation. Further, measured data accuracy increasesbecause the detection line is positioned closer to an actual location ofpaired γ-ray generation.

[0073] Next, the obtained result is reconstructed by the filtered backprojection method. The tomogram data reconstructed by the computer isstored in the storage device 12 and displayed on the display device 130.

[0074] (1) In Embodiment 1, a plurality of radiation detectors 4 arearranged in the radial direction of the through-hole 6 as well as in theaxial direction and circumferential direction thereof. Therefore, γ-raydetection signals can be derived from various radial positions of thethrough-hole 6 without reducing the amount of signal transmissionsubstance although signal transmission substance reduction could not beavoided when radiation detectors for conventional PET examination wereused. As a result, Embodiment 1 makes it possible to obtain accurateinformation about the through-hole section's radial positions reached byγ-rays (the positional information about radiation detectors 4 thatoutput γ-ray detection signals). In the conventional PET examination,one radiation detector is positioned in the radial direction of thethrough-hole 6, and a reflective material is placed in the radiationdetector to acquire the information about the through-hole section'sradial positions reached by γ-rays depending on the pattern of signaltransmission substance arrival at a photomultiplier. In such aninformation acquisition process, the signal transmission substance ispartly attenuated within the radiation detector or reflected out of theradiation detector due to the inclusion of the reflective material.Therefore, the amount of signal transmission substance decreases tolower the energy resolution.

[0075] (2) In Embodiment 1, a plurality of independent radiationdetectors 4 are arranged in the radial direction of the through-hole 6.Therefore, the entire signal transmission substance of each radiationdetector can be used for γ-ray detection to raise the radiationdetector's energy resolution. When radiation detectors having ahigh-energy resolution are used for the PET examination, γ-rays whoseenergy is attenuated by scattering can be differentiated fromunscattered, 511 keV γ-rays. As a result, an increased amount ofscattered radiation can be removed with a filter for the γ-raydiscriminator 8.

[0076] (3) Embodiment 1 provides a means of acquiring accurateinformation about the through-hole section's radial positions reached byγ-rays without decreasing the amount of signal transmission substance inradiation detectors. It is therefore possible to increase the tomogramaccuracy through the use of accurate information about γ-ray arrivalpositions and prevent the decrease in the amount of signal transmissionsubstance because the radiation detectors require no reflectivematerial. Thanks to these improvements, the energy resolution increases,thereby minimizing the influence of scattered radiation upon tomogramreconstruction. As a result, Embodiment 1 can increase the tomogramaccuracy, that is, the PET examination accuracy.

[0077] (4) Embodiment 1 uses semiconductor radiation detectors as theradiation detectors 4. Therefore, a plurality of radiation detectors 4can be arranged in the radial direction of the through-hole 6 withouthaving to increase the size of the image pickup device 2.

[0078] (5) Embodiment 1 makes it possible to easily locate a faultyradiation detector 4 in a group of radiation detectors 4 by comparingthe measured detection efficiency proportion of the radiation detectors4 with the theoretical detection efficiency proportion of the radiationdetectors 4. A faulty radiation detector 4 can easily be locatedparticularly when a plurality of radiation detectors 4 is linearlyarranged in the radial direction of the through-hole 6.

[0079] (6) In Embodiment 1, an imaging processing can be performed by asingle image pickup device 2 so as to compensate for noise caused bydetection efficiency difference and noise caused by in-vivo scatteredradiation.

[0080] In step 47 of Embodiment 1, the tomogram is reconstructed usingthe corrected PET coefficient that is obtained by correcting the countdata, which is corrected in step 45, in accordance with detectionefficiency difference in step 46. In an alternative embodiment, however,the count data corrected in step 45 can be used for tomogramreconstruction in step 47 without providing a correction in step 46.

[0081] In Embodiment 1, three radiation detectors 4 are linearlyarranged in the direction of the radius of the through-hole 6. In analternative embodiment, however, the second innermost radiation detector4 can be shifted in the circumferential direction of the through-hole 6(e.g. so that the second innermost radiation detector is positionedmidway between two adjacent innermost radiation detectors) instead ofpositioning the three radiation detectors linearly. If a plurality ofradiation detectors 4 are not linearly arranged in the direction of theradius of the through-hole 6, however, the rate of γ-ray attenuationbefore γ-ray arrival at the radiation detectors needs to be measured byconducting a test after completion of image pickup device assembly. Whena plurality of radiation detectors 4 are linearly arranged in thedirection of the radius of the through-hole 6 according to Embodiment 1,such a test need not be conducted because the γ-ray attenuation rates ofthe radiation detectors 4 are known.

[0082] In Embodiment 1, γ-ray in-vivo absorption corrections are made bymeans of transmission imaging. In an alternative embodiment, however, acommon PET correction technique can be used instead of making suchcorrections. An alternative in-vivo absorption correction method will bedescribed below. The medical examinee 17 is examined by a separatelyinstalled X-ray CT apparatus. X-rays passing through the medicalexaminee 17 are measured by a radiation detector in the X-ray CTapparatus. The rate of attenuation of an X-ray detection signalgenerated from the radiation detector is used to reconstruct thetomogram of the medical examinee 17 and determine the CT values atvarious positions within the body of the medical examinee 17. Thedetermined CT values are used to estimate the substance composition ateach position within the body of the medical examinee 17. The estimatedsubstance composition data is used to estimate the coefficient of 511keV linear attenuation at each position. The estimated linearattenuation coefficient data is used by the forward projection method todetermine the coefficient of linear attenuation between a pair ofsemiconductor devices that has detected paired γ-rays during PETexamination. The reciprocal of the determined linear attenuationcoefficient is multiplied by γ-ray detection signal count data tocompensate for a data differential arising out of attenuation within thebody. When the alternative in-vivo absorption correction methoddescribed above is used, the calibrated radiation source 31 need not beused.

[0083] Embodiment 2

[0084] A radiological imaging apparatus of another embodiment(Embodiment 2) of the present invention will be described with referenceto FIG. 8. The radiological imaging apparatus 1A of Embodiment 2 is usedfor SPECT examination. This apparatus 1A comprises an image pickupdevice 2A in place of the image pickup device 2 for the radiologicalimaging apparatus 1 and a signal processor 7A instead of the signalprocessor 7 for the radiological imaging apparatus 1. The othercomponents of the radiological imaging apparatus 1A are the same as forthe radiological imaging apparatus 1. The signal processor 7A includes aγ-ray discriminator 8A and a counter 36 connected to the γ-raydiscriminator 8A, and is provided for each radiation detector 4. Theγ-ray discriminator 8A has a filter energy setting of 120 keV althoughthe filter energy setting for the γ-ray discriminator 8 in Embodiment 1is 400 keV. The image pickup device 2A differs from the image pickupdevice 2 in that a collimator 27 is added to the former. Further, theimage pickup device 2A uses a calibrated radiation sourcecircumferential transfer unit 37A in place of the calibrated radiationsource circumferential transfer unit 37 that is used for the imagepickup device 2. The collimator 27 is positioned inside an innermostradiation detector 4 and mounted on a radiation detector support plate5. The collimator 27, which is circular, absorbs γ-rays that are aboutto obliquely fall on a radiation detector 4. As is the case withEmbodiment 1, Embodiment 2 has a plurality of radiation detector groups.

[0085] The calibrated radiation source circumferential transfer unit 37Ais provided with a guide rail 28 and a calibrated radiation sourcedevice 29A. The calibrated radiation source device 29A includes acalibrated radiation source drive 30, a calibrated radiation source 31A,and an arm 38. The calibrated radiation source 31A is mounted on the arm38. The calibrated radiation source 31A shown in FIGS. 9A and 9B uses aγ-ray source 33A in place of the γ-ray source 33, which is used for thecalibrated radiation source 31, and does not include the collimator 39,which is provided for the calibrated radiation source 31. A radiationsource emitting γ-rays of about 141 keV is used as the γ-ray source 33A.For example, 120 keV ⁵⁷CO is used.

[0086] First of all, transmission data imaging with the radiologicalimaging apparatus 1A is performed in the same manner as with theradiological imaging apparatus 1 to determine the rate of γ-rayattenuation within the body of a medical examinee. The determined γ-rayattenuation rate is used to compensate for in-vivo scattering duringSPECT examination. In Embodiment 2, the radiation detector 4 has anenergy of 120 keV.

[0087] When the emitted γ-ray energy is uniform, the γ-ray detectionefficiency can be determined from a theoretical formula. Since theradiation detectors 4 are semiconductor radiation detectors having adetection unit made of 5 mm thick CdTe, the detection efficiency of 141keV γ-rays is about 80%. Although Embodiment 2 uses a 120 keV radiationsource, appropriate results will be obtained when processing isperformed at a setting of 141 keV because the resulting detectionefficiency change is insignificant. In a radiation detector group ofthree layered radiation detectors 4, therefore, incident γ-radiation isattenuated by about 80% in the first layer radiation detector 4, and the20% γ-radiation passing through the first layer radiation detector 4 isattenuated by about 80% in the second layer radiation detector 4, thatis, about 16% γ-radiation attenuation occurs in the second layerradiation detector 4. In the third layer radiation detector 4, the 4%γ-radiation passing through the second layer radiation detector 4 isattenuated by about 80%, that is, about 3.2% γ-radiation attenuationoccurs. γ-ray detection signals reflecting such attenuations are outputfrom the radiation detectors 4. When the γ-ray detection signals fedfrom the layered radiation detectors 4 are independently measured andthe measurement result significantly differs (by, for instance, morethan ±5%) from the theoretical detection efficiency proportion(80:16:3.2) of the layered radiation detectors 4, it means that one ormore of the radiation detectors 4 is deteriorated. As is the case withEmbodiment 1, Embodiment 2 also makes it possible to locate adeteriorated radiation detector 4 within a radiation detector group anddetermine the percentage of detection efficiency decrease caused bydeterioration. The concept of fault detection provided by Embodiment 2will be described below.

[0088] In Embodiment 2, γ-rays emitted at a certain time from the γ-raysource 33A fall on three radiation detectors 4 in a radiation detectorgroup due to the shape of the collimator 27, but do not fall on threeradiation detectors in a radiation detector group adjacent to the aboveradiation detector group. However, the concept of fault detectiondescribed with reference to Embodiment 1 is still applied to Embodiment2, and the process applied to Embodiment 2, which will be describedlater, is virtually the same as that is indicated in FIGS. 5 and 6.

[0089] A SPECT examination process performed with the radiologicalimaging apparatus 1A will be described below. A SPECT pharmaceutical isadministered to a medical examinee 17. A bed 16 on which the medicalexaminee 17 is laid is inserted into a through-hole 6 of an image pickupdevice 2A. The SPECT pharmaceutical is gathered at a lesion in themedical examinee 17. The lesion in the body of the medical examinee 17emits 141 keV γ-rays (when the SPECT pharmaceutical contains ⁹⁹Tc).Radiation detectors 4 detect the emitted γ-rays as is the case withEmbodiment 1. The γ-ray detection signals output from the radiationdetectors 4 are received by the associated γ-ray discriminator 8A. Theγ-ray discriminator 8A uses a filter to pass a γ-ray detection signal(excluding scattered γ-rays) having energy higher than an energy settingof 120 keV, and generates a pulse signal having an appropriate energyfor the γ-ray detection signal. A counter 36 uses the pulse signal toperform counting and determines the count data for the γ-ray detectionsignal. The counter 36 outputs the count data and the positionalinformation about detection points (positional information about aradiation detector 4 that generated the γ-ray detection signal). Acomputer 11 associates the count data with the positional informationabout detection points and stores it in a storage device 12.

[0090] The computer 11 uses the count data and other relevant data toperform a processing procedure indicated in FIGS. 5 and 6 to reconstructa tomogram of the medical examinee 17. Embodiment 2 is different fromEmbodiment 1 in processing steps 40, 41, 44, 46, and 47, but is equal toEmbodiment 2 in the other processing steps. Therefore, the descriptionof Embodiment 2 will cover processing steps 40, 41, 44, 46, and 47 only.In step 40 of Embodiment 2, the count data derived from SPECTexamination, the positional information about the associated detectionpoints, and the count data derived from transmission imaging are readfrom the storage device 12 and input. In step 41, the theoreticaldetection efficiency proportion of grouped radiation detectors iscalculated. This theoretical proportion can be determined by performingtheoretical calculations on the transmission distance of γ-rays emittedfrom the medical examinee 17 (this distance varies with theradionuclides contained in the SPECT pharmaceutical) and the sequence ofγ-ray transmission. When a SPECT pharmaceutical containing ⁹⁹Tc isadministered to the medical examinee 17, the theoretical detectionefficiency proportion of the first-layer to third-layer radiationdetectors 4 in a radiation detector group is about 80:16:3.2.

[0091] In step 44 of Embodiment 2, the attenuation correction count forradiation detectors is calculated. Although paired γ-rays is emittedfrom a lesion during PET examination, a single γ-ray is emitted duringthe SPECT examination. Therefore, step 44 of Embodiment 2 differs fromthe counterpart of Embodiment 1. The attenuation correction count forradiation detectors in a radiation detector group is calculated. Asregards the γ-radiation emission from a lesion during SPECT examination,one count data derived from SPECT examination and the γ-ray attenuationrate calculated in step 43 are used to reconstruct the tomogram of themedical examinee 17. First, the transmission image obtained in step 43is back-projected to determine the γ-ray attenuation rate of eachposition within the body of the medical examinee 17. The determinedγ-ray attenuation rate is used to estimate the substance composition ateach position within the body of the medical examinee 17. The estimatedsubstance composition data is used to estimate the coefficient of 141keV linear attenuation at each position within the body. The estimatedlinear attenuation coefficient data is used by the forward projectionmethod to determine the average linear attenuation coefficient for caseswhere γ-rays are generated so that they are incident on a certainradiation detector via the collimator 27. The reciprocal of thedetermined linear attenuation coefficient is the attenuation correctioncount.

[0092] In step 46 of Embodiment 2, the γ-ray image pickup signal iscorrected in accordance with the detection efficiency difference amongradiation detectors. In SPECT examination where a single γ-ray isemitted, the count data is corrected using the detection efficiency of aradiation detector group at which the single γ-ray arrives. Thiscorrection is made in accordance with equation (2), which is describedwith reference to Embodiment 1. The value Xsi_(ij) in equation (2) iscorrected SPECT count data. The corrected SPECT count data calculatedfrom equation (2) is stored in the storage device 12. In step 47, thetomogram is reconstructed using the corrected SPECT count data Xsi_(ij),which is derived from the correction in step 46, and the positionalinformation about detection points.

[0093] Embodiment 2 also provides advantages (1) through (6) ofEmbodiment 1.

[0094] Embodiment 3

[0095] A radiological imaging apparatus of another embodiment(Embodiment 3) of the present invention will be described with referenceto FIGS. 10 and 11. The radiological imaging apparatus 1B of Embodiment3 is used for X-ray CT examination (in which an X-ray emission from anX-ray source 60 passes through the body of a medical examinee and isdetected by radiation detectors) and PET examination. This apparatus 1Bcomprises an image pickup device 2B in place of the image pickup device2 for the radiological imaging apparatus 1 and a signal processor 7Ainstead of the signal processor 7 for the radiological imaging apparatus1. The other components of the radiological imaging apparatus 1B are thesame as for the radiological imaging apparatus 1. The image pickupdevice 2B uses a calibrated radiation source circumferential transferunit 37B in place of the calibrated radiation source circumferentialtransfer unit 37 that is used for the image pickup device 2. Thecalibrated radiation source circumferential transfer unit 37B includes aguide rail 28 and a calibrated radiation source device 29B. Thecalibrated radiation source device 29B includes a calibrated radiationsource drive 30, a calibrated radiation source 31, an X-ray source 60,and an arm 38. The calibration radiation source 31 and X-ray source 60are mounted on the end of the arm 38. The calibration radiation source31 and X-ray source 60 may be mounted on the end of the arm 38 so thatthey are aligned in the circumferential direction of a through-hole 6.The calibrated radiation source circumferential transfer unit 37Bdoubles as an X-ray source circumferential transfer unit. The calibratedradiation source drive 30 doubles as an X-ray source drive. Embodiment 3includes a drive controller 35 and a radiation source controller 69.

[0096] The X-ray source 60 includes a publicly known X-ray tube, whichis not shown. This X-ray tube is provided with an anode, a cathode, acurrent source for the cathode, and a voltage source for applying avoltage between the anode and cathode, which are mounted inside anexternal cylinder. The cathode is formed of a tungsten filament.Electrons are emitted from the filament when a current flows from thecurrent source to the cathode. These electrons are accelerated by avoltage (several hundred kV) applied from the voltage source between thecathode and the anode, and collide with the anode (W, Mo, etc.), whichis the target. Collision of electrons with the anode produces X-rays of80 keV. These X-rays are emitted from the X-ray source 60.

[0097] The signal processor 7A includes a signal discriminator 61, aγ-ray discriminator 8 not included in the signal discriminator 61, and acoincidence counter 9. The signal discriminator 61 is connected to eachof the first-layer (4X) radiation detectors 4, which are the innermostradiation detectors in each radiation detector group. As shown in FIG.12, each of the signal discriminators 61 includes a selector switch 62,a γ-ray discriminator 8, and an X-ray signal processor 66. The selectorswitch 62 includes a movable terminal 63 and stationary terminals 64 and65. The first-layer (4X) radiation detectors 4 are connected to themovable terminal 63 on the selector switch 62 via a wiring 13. The γ-raydiscriminator 8 is connected to the stationary terminal 64, and theX-ray signal processor 66 is connected to the stationary terminal 65.The minus terminal on a power supply 68 is connected to the wiring 13via a resistor 67. The plus terminal on the power supply 65 is connectedto radiation detectors 4. The radiation detectors in the secondinnermost layer (4Y) and third innermost layer (4Z) within eachradiation detector group are connected to their respective γ-raydiscriminators 8 as is the case with Embodiment 1. All the γ-raydiscriminators 8, including the γ-ray discriminator 8 within the signaldiscriminator 61, are connected to one coincidence counter 9. Thecoincidence counter 9 may be furnished for each division of radiationdetectors 4 as is the case with Embodiment 1. The coincidence counter 9and X-ray signal processor 66 are connected to the computer 11.

[0098] First, a transmission data imaging process is performed with thecalibrated radiation source 31 as is the case with Embodiment 1. Aftercompletion of transmission data imaging, the image pickup device 2B isused to conduct a radiological examination (PET examination and X-ray CTexamination).

[0099] An X-ray CT examination/PET examination process performedaccording to Embodiment 3 will be described below. A PET pharmaceuticalis injected into or otherwise administered to a medical examinee 17 inadvance in such a manner that the radioactivity administered to the bodyof the medical examinee is 370 MBq. When a predetermined period of timeelapses, a bed 16, on which the medical examinee 17 is laid, is insertedinto the through-hole 6 of the image pickup device 2B so as to positionthe medical examinee 17 within the through-hole 6. An X-ray CTexamination/PET examination is conducted with the image pickup device2B.

[0100] Before giving a detailed description of radiological examinationaccording to Embodiment 3, the principles of radiation detectionprovided by Embodiment 3 will be described. X-rays emitted from theX-ray source are oriented in a specific direction and incident on themedical examinee for a predetermined period of time, and X-rays passingthrough the body of the medical examinee are detected by radiationdetectors. This X-ray detection operation (scan) is repeated. The datafor an X-ray CT image (tomogram that is derived from X-ray CT andcontains an image of a medical examinee's internal organs and bones) iscreated in accordance with the intensity of X-rays detected by aplurality of radiation detectors. For the acquisition of highly accurateX-ray CT image data, it is desirable that γ-rays emitted from theinterior of the medical examinee's body due to the administered PETpharmaceutical be not incident on the radiation detectors that aredetecting X-rays during X-ray CT examination. The influence of γ-rays ona single radiation detector is negligible when the duration of a medicalexaminee's exposure to X-rays is shortened in accordance with the rateof γ-ray incidence. Therefore, efforts have been made to reduce theperiod of time during which the medical examinee is exposed to X-rays.The rate of γ-ray incidence on a radiation detector is considered todetermine the duration of X-ray exposure T. When the in-vivoradioactivity based on a PET pharmaceutical to be administered to amedical examinee for PET examination is N (Bq), the rate of generatedγ-ray passage through the body of the medical examinee is A, the rate ofincidence determined from a solid angle of a radiation detector is B,and the sensitivity of a detection device is C, the rate of γ-raydetection by the radiation detector α (counts/sec) is given by equation(3). In equation (3), the coefficient “2” means that a pair of γ-rays(two γ-rays) are emitted when a positron annihilates.

α=2NABC  (3)

[0101] The probability W with which γ-rays are detected by a detectiondevice within the irradiation time T is given by equation (4).

W=1−exp(−Tα)  (4)

[0102] When the irradiation time T is determined so as to minimize thevalue W, the influence of γ-rays incident on a radiation detector duringX-ray CT examination is negligible.

[0103] A typical X-ray irradiation time T will be described below. TheX-ray irradiation time T is determined from equations (3) and (4). Themaximum radiation intensity prevailing within the body of a medicalexaminee due to a PET pharmaceutical administered to the medicalexaminee for PET examination is about 370 MBq (N=370 MBq). The rate ofγ-ray passage through the medical examinee's body A is about 0.6 (A=0.6)on the presumption that the medical examinee's body is water having aradius of 15 cm. If, for instance, 5 mm square radiation detectors arearranged in the form of a ring having a radius of 50 cm, the rate ofincidence B determined from the solid angle of one radiation detector is8×10⁻⁶ (B=8×10⁻⁶). The radiation detector's maximum detectionsensitivity C is about 0.6 (C=0.6) when semiconductor radiationdetectors are used. These values indicate that the rate of γ-raydetection by one radiation detector α is about 2000 counts/second. If,for instance, the X-ray irradiation time T is 1.5 μsec, the probabilityW with which a radiation detector detects γ-rays during an X-raydetection process is 0.003. It means that such γ-rays are practicallynegligible. If the X-ray irradiation time is 1.5 μsec or shorter insituations where the radioactivity administered to the body is 360 MBqor less, W<0.003. It means that the γ-ray detection probability is 0.3%or lower and negligible.

[0104] An X-ray CT examination/PET examination performed with the imagepickup device 2B according to the above principles will be detailedbelow.

[0105] To start an X-ray CT examination, the drive controller 35 outputsa drive start signal to close a switch (hereinafter referred Lo as themotor switch) that is connected to a motor for the calibrated radiationsource drive 30 and to a power supply. An electrical current is suppliedto rotate the motor. The turning force of the motor is transmitted to apinion via a speed reduction mechanism so that the calibrated radiationsource device 29B, that is, the X-ray source 60 circumferentially movesalong the guide rail 28. The X-ray source 60 moves around the medicalexaminee 17 at a preselected speed while it is positioned within thethrough-hole 6. At the end of the X-ray CT examination, the drivecontroller 35 outputs a drive stop signal to open the motor switch. Thisstops the movement of the X-ray source 60 in the circumferentialdirection. In Embodiment 3, the radiation detectors 4, which arearranged circumferentially in a circular form, do not move in thecircumferential direction or in the axial direction of the through-hole6. For a control signal transmission from the immobile X-ray sourcecontroller and drive controller to the mobile X-ray source device, apublicly known technology that does not obstruct the movement of theX-ray source device is used.

[0106] The radiation source controller 69 controls the time of X-rayemission from the X-ray source 60. More specifically, the radiationsource controller 69 repeatedly outputs an X-ray generation signal andX-ray shut-off signal. The first X-ray generation signal is output inaccordance with the input of the above drive start signal to theradiation source controller 69. Upon X-ray generation signal output, aswitch (this switch is hereinafter referred to as the X-ray sourceswitch; not shown) provided between the X-ray tube anode (or cathode) ofthe X-ray source 60 and power supply closes. When a first preselectedperiod of time elapses, the X-ray shut-off signal is output to open theX-ray source switch. When a second preselected period of time elapses,the X-ray source switch closes. Radiation source control is repeatedlyexercised in this manner. For the first preselected period of time, avoltage is applied between the anode and cathode. For the secondpreselected period of time, however, such a voltage application does nottake place. Thanks to the control exercised by the radiation sourcecontroller 69, the X-ray tube emits 80 keV X-rays in a pulsating manner.The irradiation time T, which is the first preselected period of time,is set, for instance, to 1 μsec so that the γ-ray detection probabilityat radiation detectors 4 can be neglected. The second preselected periodof time is the time interval T0 during which the X-ray source 60 movesfrom one radiation detector 4 to a circumferentially adjacent radiationdetector 4, and determined by the speed at which the X-ray source 60moves circumferentially on the guide rail 28. The first and secondpreselected periods of time are stored in the radiation sourcecontroller 69.

[0107] When the X-ray shut-off signal and X-ray generation signal arerepeatedly output, the X-ray source 60 emits X-rays for the firstpreselected period of time, that is, 1 μsec, and halts its X-rayemission for the second preselected period of time. This X-ray emissionand shut-off cycle is repeated while the X-ray source 60 moves in thecircumferential direction.

[0108] X-rays emitted from the X-ray source 60 fall on the medicalexaminee 17 in the form of a fan beam. As the X-ray source 60 moves inthe circumferential direction, X-rays come from the circumference tofall on the medical examinee 17. X-rays passing through the medicalexaminee 17 are detected by a plurality of radiation detectors 4circumferentially positioned around a radiation detector 4 that ismounted 180 degrees away from the X-ray source 60 when the axial centerof the through-hole 6 is regarded as the base point. These radiationdetectors output the detection signals related to the detected X-rays.The X-ray detection signals are then entered in the respective signaldiscriminators 61 via the associated wirings 13. These X-ray detectingradiation detectors 4 are referred to as first radiation detectors 4 forthe sake of convenience.

[0109] From the medical examinee 17 on the bed 16, 511 keV γ-rays areemitted due to the administered PET pharmaceutical. Radiation detectors4 other than the first radiation detectors 4 output γ-ray detectionsignals. These γ-ray detecting radiation detectors 4 are referred to assecond radiation detectors 4 for the sake of convenience. The γ-raydetection signals output from the second radiation detectors in thefirst layer are delivered to the respective signal discriminators 61 viathe associated wirings 13. The γ-ray detection signals output from thesecond radiation detectors in the second and third layers are deliveredto the respective γ-ray discriminators 8 via the wirings 13. Note thatonly the radiation detectors 4 in the first layer are connected to thesignal discriminators 61. The reason is that almost all X-rays (morethan 90%) passing through the medical examinee 17 are detected by theradiation detectors 4 in the first layer since the X-ray energy is 80keV.

[0110] Within the signal discriminator 61, the γ-ray detection signaloutput from a second radiation detector 4 in the first layer is conveyedto a γ-ray discriminator 8, and the X-ray detection signal output from afirst radiation detector 4 is conveyed to the X-ray signal processor 66.These detection signal transmission operations are performed inaccordance with a switching operation of the selector switch 62 of thesignal discriminator 61. The switching operation for connecting themovable terminal 63 of the selector switch to the stationary terminal 64or 65 is performed in accordance with a switching control signal that isoutput from the drive controller 35. The drive controller 35 selects thefirst radiation detector 4 from the radiation detectors 4 in the firstlayer, and connects the movable terminal 63 to the stationary terminal65 in the signal discriminator 61 to be connected to the first radiationdetector 4. The theoretical detection efficiency proportion of threelayered radiation detectors in a radiation detector group is 20:16:12.8(the values are arranged in order from the innermost detector to theoutermost).

[0111] The selection of the first radiation detectors 4 will bedescribed. An encoder (not shown) is linked to a motor in the calibratedradiation source drive 30. The drive controller 35 inputs the encoder'sdetection signal, determines the circumferential position of thecalibrated radiation source drive 30, that is, the X-ray source 60, anduses the stored positional data about radiation detectors 4 to select aradiation detector 4 that is positioned 180° away from the X-ray source60. Since the X-rays emitted from the X-ray source 60 has a width in thecircumferential direction of the guide rail 28, not only the selectedradiation detector 4 but also the other radiation detectors positionedin the circumferential direction detect X-rays passing through themedical examinee 17. The drive controller 35 selects such additionalradiation detectors as well. These radiation detectors are the firstradiation detectors. As the X-ray source moves in the circumferentialdirection, the first radiation detectors 4 change. It looks as if thefirst radiation detectors 4 moved in the circumferential directionduring the circumferential travel of the X-ray source 60. When the drivecontroller 35 selects another radiation detector 4 during thecircumferential travel of the X-ray source 60, the movable terminal 63connected to the newly designated first radiation detector 4 isconnected to stationary terminal 65. When the movable terminal 63 isconnected to a radiation detector 4 that is no longer the firstradiation detector 4 due to the circumferential travel of the X-raysource 60, it is connected to the stationary terminal 64 by the drivecontroller 35. A radiation detector in the first layer becomes a firstradiation detector 4 at a certain time and becomes a second radiationdetector 4 at another time, depending on the positional relationship tothe X-ray source 60. Therefore, a radiation detector 4 in the firstlayer outputs both an X-ray image pickup signal and a γ-ray image pickupsignal at different times.

[0112] A first radiation detector 4 detects X-rays passing through themedical examinee 17 after being emitted from the X-ray source 60 for afirst preselected period of time, that is, 1 μsec. The probability withwhich the first radiation detector 4 detects a γ-ray emission from themedical examinee 17 for a 1 μsec period is very low and negligible asexplained earlier. Many γ-rays generated within the body of the medicalexaminee 17 due to the administered PET pharmaceutical are not emittedin a specific direction but emitted in all directions. As describedearlier, these γ-rays are paired, emitted in almost opposite directions(180°±0.6°), and detected by a certain second radiation detector 4.

[0113] A signal process performed by a signal discriminator 61 when itreceives an X-ray detection signal/γ-ray detection signal output from aradiation detector 4 in the first layer will be described. As explainedearlier, the X-ray detection signal output from a first radiationdetector 4 is received by the X-ray signal processor 66 via the selectorswitch 62. The X-ray signal processor 66 uses an integrator to performcalculations on the input X-ray detection signal and outputs theinformation about the integrated X-ray detection signal value, that is,the measured X-ray intensity. The intensity information about the X-raydetection signal is conveyed to the computer 11 and stored in thestorage device 12 by the computer 11. The γ-ray detection signal outputfrom a second radiation detector 4 in the first layer is received by aγ-ray discriminator 8 via the selector switch 62. The γ-raydiscriminator 8 for a signal discriminator 61 generates a pulse signalhaving a predetermined energy when it receives a γ-ray detection signalhaving an energy greater than an energy setting (400 keV). As is thecase with Embodiment 1, a coincidence counter 9 receives pulse signalsoutput from all γ-ray discriminators 8, and outputs the count data abouteach γ-ray detection signal and the positional information about the twopoints of paired γ-ray detection. The count data and positionalinformation are conveyed to the computer 11 and stored in the storagedevice 12 by the computer 11.

[0114] The computer 11 performs a process indicated in FIG. 13. TheX-ray detection signal intensity, the count data and positionalinformation about the associated detection points derived from PETexamination, and the count data derived from transmission data imagingare read from the storage device 12 and input (step 69). The rate ofX-ray attenuation in each voxel within the body of the medical examinee17 is calculated from the X-ray detection signal intensity (step 70).The calculated X-ray attenuation rate is stored in the storage device12. The tomogram of the medical examinee 17 is reconstructed using therate of X-ray detection signal attenuation at the associated positions(step 71). The tomogram reconstructed using the X-ray detection signalattenuation rate is referred to as an X-ray CT image. For X-ray CT imagereconstruction purposes, the X-ray detection signal attenuation rateread from the storage device 12 is used to determine the coefficient oflinear attenuation within the body of the medical examinee 17 betweenthe X-ray source 60 and the semiconductor device unit of a firstradiation detector 4. This linear attenuation coefficient is used todetermine the linear attenuation coefficient of each voxel by thefiltered back projection method. The linear attenuation coefficient ofeach voxel is used to determine the CT value of each voxel. Thedetermined voxel CT values are used to obtain X-ray CT image data. TheX-ray CT image data is stored in the storage device 12. Next, thecross-sectional tomogram of the medical examinee 17 is reconstructedusing the γ-ray detection signal count data about the associatedpositions and the positional information about detection points (step72). The tomogram reconstructed using the γ-ray detection signal countdata is referred to as a PET image. In step 72, processing steps 41through 47 in FIG. 5, which is used for the description of Embodiment 1,are performed to obtain a PET image. The obtained PET image data isstored in the storage device 12. The PET image data and X-ray CT imagedata are synthesized to obtain synthesized tomogram data, which containsboth the PET image data and X-ray CT image data. The resultingsynthesized tomogram data is stored in the storage device 12 (step 73).Synthesis of PET image data and X-ray CT image data can be achievedeasily and accurately by aligning a reference point common to these twoimage data (e.g., central axis position of the through-hole 6).Positional alignment can be accurately achieved because the PET imagedata and X-ray CT image data are generated according to detectionsignals output from shared radiation detectors 4. The synthesizedtomogram data is recalled from the storage device 12, output to thedisplay device 130 (step 74), and displayed on the display device 130.Since the synthesized tomogram displayed on the display device 130contains an X-ray CT image, a diseased area visualized by a PET imagecan easily be located within the body of a medical examinee 17. Morespecifically, since the X-ray CT image contains an image of internalorgans and bones, doctors can locate a diseased area (e.g., cancerousarea) based on the relationship to the internal organs and bones.

[0115] Embodiment 3 provides the following advantages in addition toadvantages (1) through (6) of Embodiment 1.

[0116] (7) In Embodiment 3, it is possible to detect not only aplurality of paired γ-rays emitted from a medical examinee 17 or asubject with radiation detectors 4 arranged around the circumference ofthe through-hole 6 but also X-rays that are emitted from acircumferentially moving X-ray source 60 and passed through the medicalexaminee 17 (with radiation detectors 4 in the first layer). Although aconventional technology required the use of an image pickup device fordetecting transmitted X-rays and the use of another image pickup devicefor detecting γ-rays, Embodiment 3 requires the use of only one imagepickup device and simplifies the structure of a radiological imagingapparatus that provides both X-ray CT examinations and PET examinations.

[0117] (8) In Embodiment 3, each of first-layer radiation detectors 4arranged around the circumference of the through-hole 6 outputs both anX-ray detection signal and γ-ray detection signal. This configurationcontributes toward radiological imaging apparatus structuresimplification and downsizing.

[0118] (9) Embodiment 3 uses X-ray detection signals output fromradiation detectors 4 in the first layer to reconstruct a first tomogram(X-ray CT image) of a medical examinee 17, which contains an image ofinternal organs and bones. It also uses γ-ray detection signals outputfrom radiation detectors 4 in the first to third layers to reconstruct asecond tomogram (PET image) of the medical examinee 17, which containsan image of a diseased area. Since first and second tomogram data arereconstructed in accordance with the signals output from radiationdetectors 4 that are mounted around the circumference of thethrough-hole 6 in an image pickup device 2B, the first and secondtomogram data can be synthesized with their positional relationshipaccurately adjusted. Therefore, an accurate tomogram (synthesizedtomogram) containing an image of a diseased area, internal organs, andbones can be obtained with ease. The resulting synthesized tomogrammakes it possible to accurately locate a diseased area based on therelationship to the internal organs and bones. The first and secondtomogram data can easily be synthesized by, for instance, effectingtomogram alignment with respect to the axial center of the through-hole6 in the image pickup device 2B.

[0119] (10) In Embodiment 3, detection signals necessary for thecreation of a first tomogram and detection signals necessary for thecreation of a second tomogram can be obtained from shared radiationdetectors 4. Therefore, the time required for the examination of amedical examinee 17 (examination time) can be considerably reduced. Inother words, the detection signals necessary for the creation of thefirst tomogram and the detection signals necessary for the creation ofthe second tomogram can be obtained within a short examination time.Unlike a conventional technology, Embodiment 3 minimizes the probabilitywith which a medical examinee moves because it does not have to transfera medical examinee from an image pickup device for transmitted X-raydetection to another image pickup device for γ-ray detection. Since thenecessity for transferring a medical examinee from an image pickupdevice for transmitted X-ray detection to another image pickup devicefor γ-ray detection is eliminated, the time required for the examinationof a medical examinee can be reduced.

[0120] (11) Since the amount of γ-ray image pickup signal input to theX-ray signal processor 66, that is, a first signal processor isconsiderably reduced, accurate data about a first tomogram can beobtained. Therefore, when image data derived from the synthesis of thedata about a first tomogram and the data about a second tomogram isused, diseased areas can be located with increased accuracy.

[0121] (12) In Embodiment 3, the X-ray source 60 circulates inside anumber of arrayed radiation detectors. Therefore, the diameter of thethrough-hole 6 increases, making it possible to increase the number ofradiation detectors 4 to be mounted in the first layer. Increasing thenumber of radiation detectors 4 mounted in the circumferential directionenhances the sensitivity and improves the cross-sectional imageresolution of a medical examinee 17.

[0122] (13) In Embodiment 3, the arm 38 on which the X-ray source 60 ismounted and the X-ray source 60 are positioned inside radiationdetectors 4. Therefore, they could obstruct the γ-rays emitted from amedical examinee 17 and prevent radiation detectors 4 positionedimmediately behind them from detecting such γ-rays, resulting in theloss of detected data necessary for PET image formation. In Embodiment3, however, the calibrated radiation source drive 30 rotates the X-raysource 60 and arm 38 in the circumferential direction as describedearlier. Therefore, Embodiment 3 does not incur any substantial loss ofdata. It should be noted in this connection that the X-ray source 60 andarm 38 rotate at a rate of about 1 second per slice. The time requiredfor the rotation of the X-ray source and arm is considerably shorterthan the minimum time required for PET examination, which is on theorder of several minutes. It means that no substantial data loss canoccur.

[0123] Embodiment 4

[0124] A radiological imaging apparatus of another embodiment(Embodiment 4) of the present invention will be described with referenceto FIG. 14. The radiological imaging apparatus 75 of Embodiment 4 is adigital X-ray examination apparatus that uses a flat panel detector. Theradiological imaging apparatus 75 includes an X-ray source 76 that issupported by a stanchion 77, a flat panel detector provided with aplurality of radiation detectors (not shown) and supported by astanchion 79, X-ray signal processors 66, and an X-ray imaging device80. Within the flat panel detector 78, a large number of radiationdetectors 4 are arranged in the direction of the height and in thedirection of the width. As indicated by radiation detectors 4 i, 4 j,and 4 k in FIG. 15, the radiation detectors 4 are also linearly arrangedin the direction of the depth (in the traveling direction of X-rayspassing through a medical examinee 17) so as to form three layers ofradiation detectors. A plane 82 shown in FIG. 15 faces the X-ray source76. The X-ray signal processor 66 is connected to the radiationdetectors 4. The X-ray imaging device 80 includes a computer 11, astorage device 12, and a display device 130. The storage device 12 anddisplay device 130 are connected to the computer 11 to which all theradiation detectors are connected.

[0125] An X-ray examination performed with the radiological imagingapparatus 75 will be described below. A medical examinee 17 standsbetween the X-ray source 76 and flat panel detector 78 with its backfacing the X-ray source 76. X-rays emitted from the X-ray source 76 passthrough the medical examinee 17 and are detected by the radiationdetectors 4 in the flat panel detector 78. The radiation detectors 4detect X-rays and output X-ray detection signals. The X-ray signalprocessors 66 add up the X-ray detection signals so as to output theinformation about X-ray intensity. The X-ray intensity informationoutput from each X-ray signal processor 66 is entered in the computer 11and stored in the storage device 12. The computer 11 acquires the X-rayintensity information from the storage device 12 and calculates therates of X-ray attenuation at various positions within the body of themedical examinee 17.

[0126] A radiation detector group is formed by three layered radiationdetectors 4 that are linearly arranged in the direction of the depth,beginning with the plane 82 of the flat panel detector 78. In Embodiment4, the detection efficiency proportion of radiation detectors 4 in aradiation detector group also varies with the energy of X-rays emittedfrom the X-ray source 76. When, for instance, the radiation detectors 4having a detection unit, which is a 2 mm cube made of CdTe, detect 100keV rays passing through a medical examinee 17, the theoreticaldetection efficiency proportion prevailing within the radiation detectorgroup is about 84:13:2.5. This theoretical detection efficiencyproportion is stored in the storage device 12.

[0127] The computer 11 uses the X-ray intensity information stored inthe storage device 12 to calculate the measured detection efficiencyproportion of radiation detectors 4 in each radiation detector group.The computer 11 performs processing step 52 of Embodiment 1. When thedeviation of the calculated measured detection efficiency proportionsfrom the theoretical one is within a predefined range, the computer 11uses the above-calculated X-ray attenuation rate to generate grayscaleimage data for X-ray imaging of the medical examinee 17. If theabove-mentioned deviation is outside the predefined range, on the otherhand, the computer 11 performs processing steps 53, 54, and 55 ofEmbodiment 1. When the detection efficiency of a deteriorated radiationdetector is corrected in processing step 55, the X-ray intensity forthat radiation detector is determined according to the correcteddetection efficiency, and the above-mentioned X-ray attenuation rate iscorrected with the determined X-ray intensity taken into account. Thecomputer 11 uses the corrected X-ray attenuation rate to generate theabove-mentioned grayscale image data.

[0128] Embodiment 4 provides advantages (1) through (5) of Embodiment 1.However, it should be noted that advantage (3) results in an increase inthe X-ray image accuracy.

[0129] In Embodiment 4, the radiation detectors do not always have to belinearly arranged in the direction of the depth of the flat paneldetector 78. They can be alternatively arranged so that all theradiation detectors 4 in the second layer overlap with two radiationdetectors in the first layer (as viewed from the plane 82).

[0130] Next, the correction method to be used when the radiationdetectors are not linearly arranged will be described with reference toan example in which a flat panel detector is used for digital X-rayexamination. Although a radiological imaging apparatus having a flatpanel detector is the same as indicated in FIG. 10, the first- tothird-layer radiation detectors 4 for the flat panel detector 70 arenonlinearly arranged with the second-layer radiation detectors 4 shownin FIG. 15 displaced laterally. Even when the radiation detectors 4 arearranged in this manner to form multiple layers, it is possible tolocate part of deteriorated radiation detectors 4 and providecorrections for the measured values of deteriorated radiation detectors.

[0131] Embodiment 5

[0132] A radiological imaging apparatus of another embodiment(Embodiment 5) of the present invention will be described with referenceto FIG. 16. The radiological imaging apparatus 83 of Embodiment 5 is anX-ray CT apparatus. The radiological imaging apparatus 83 includes anX-ray source 84 that is mounted on an arm 86, a radiation detector unit85 mounted on the arm 86, X-ray signal processors 66, and a tomogramgenerator 88. The arm 86 is supported by a stanchion 87. The X-raysource 84 and radiation detector unit 85 face each other and arepositioned away from each other so that a medical examinee 17 can bepositioned between them. As is the case with the flat panel detector 78,the radiation detector unit 85 is equipped with a large number ofradiation detectors 4. The radiation detectors 4 are not only arrangedin the direction of the height and in the direction of the width, butalso linearly arranged in the direction of the depth, beginning with theplane facing the X-ray source 84, so as to form three radiationdetection layers. The arm 86 can be rotated, although the details of itsmechanism are not shown, so that the X-ray source 84 and radiationdetector unit 85 move around the medical examinee 17 lying on a bed 16.

[0133] An examination performed with the radiological imaging apparatus83 will be described below. A medical examinee 17 lying on the bed ispositioned between the X-ray source 84 and radiation detector unit 85.X-rays emitted from the X-ray source 84 fall on the medical examinee 17and pass through the body of the medical examinee 17. The X-raystransmitted in this manner are detected by the radiation detectors 4 inthe radiation detector unit 85. The rotating device (not shown) for thearm 86 rotates the X-ray source 84 and radiation detector unit 85 aroundthe medical examinee 17 (through 180° or 360° relative to a certaincross section of the medical examinee 17). X-ray detection signalsoutput from the radiation detectors 4 are entered in the respectiveX-ray signal processors 66. The X-ray signal processors 66 determine theX-ray intensity in accordance with the measurements of the X-raydetection signals. In accordance with the X-ray intensity, a computer 11calculates the medical examinee's in-vivo X-ray attenuation rate thatprevails between the rotating X-ray source 84 and the portion of therotating radiation detector unit 85 that faces the X-ray source 84. Thelinear attenuation coefficient determined in this manner is stored inthe storage device 12.

[0134] In the same manner as in step 51 of Embodiment 1, the computer 11calculates the detection efficiency proportion of three layeredradiation detectors 4 that are linearly arranged beginning with theplane facing the X-ray source 84 for the radiation detector unit 85, andthen continues to perform processing steps 52, 53, 54, and 55 ofEmbodiment 1. When the deviation of the calculated measured detectionefficiency proportion from the theoretical one is found in processingstep 52 to be within a predefined range, the computer 11 uses theaforementioned filtered back projection method or the like to determinethe linear attenuation coefficient of each voxel from thedetector-to-radiation source X-ray attenuation rate stored in thestorage device 12, and converts the resulting value to a CT value. Ifthe above deviation is found in processing step 52 to be outside thepredefined range, on the other hand, the computer 11 uses the correcteddetection efficiency value to correct the linear attenuation coefficientstored in the storage device 12, and calculates the CT value from thecorrected linear attenuation coefficient. The computer 11 uses the CTvalue of each voxel to reconstruct the X-ray CT image.

[0135] Embodiment 5 provides the advantages of Embodiment 4.

[0136] Embodiment 6

[0137] When γ-rays emitted from a medical examinee due to anadministered PET pharmaceutical fall on a radiation detector, theyattenuate although in some cases they may pass through as they are. Whenγ-ray attenuation occurs within a radiation detector, the radiationdetector outputs a detection signal (electrical charge) that correspondsto the γ-ray energy attenuation. Detected (attenuated) γ-rays scatterwithin the radiation detector except when they suffer total attenuation.Scattered γ-rays change their direction of travel and fall on anotherradiation detector at a different angle of incidence. It goes withoutsaying that scattered γ-rays may pass through a radiation detectorwithout suffering any subsequent attenuation, suffer total attenuationwithin another radiation detector, or scatter again and become detected.In other words, γ-rays detected by a radiation detector may be eitherunscattered γ-rays (not scattered by a radiation detector) or scatteredγ-rays.

[0138] As described above, γ-rays change the direction of their travelwhen they scatter. Therefore, the source of γ-ray generation does notexist on the extension of a scattered γ-ray vector. That is, PET imagedata based on scattered γ-ray detection signals turns out to beerroneous, causing an error. In consideration of energy attenuation uponγ-ray scattering, therefore, y-rays having an energy smaller than apredefined threshold energy value were conventionally considered to bescattered and then removed. When such a method was employed, however,unscattered γ-rays were frequently regarded as scattered γ-rays simplybecause their energy was below the above-mentioned threshold energyvalue so that the PET image data collection efficiency was lowered.

[0139] A nuclear medicine diagnostic apparatus described in JP-A No.321357/2000 subjects γ-ray detection signals to coincidence countingwhen it detects a plurality of γ-rays, concludes that nearlysimultaneously detected γ-rays are generated from the same source,checks whether the calculated total energy of the detected γ-rays iswithin a predefined range, and determines whether unscattered γ-rays areincluded in the detected γ-rays. When it concludes that unscatteredγ-rays are included, it selects an initial γ-ray incidence position bypicking up one detected γ-ray having a statistically high probability ofbeing unscattered. However, the initial incidence position determined bythis conventional technology is selected probabilistically and may be inerror. Therefore, this conventional technology merely provides a limitedincrease in the detection.

[0140] As a solution to the problem with the nuclear medicine diagnosticapparatus described in JP-A No. 321357/2000, a radiological imagingapparatus of another embodiment (Embodiment 6) of the present inventionwill be described with reference to FIGS. 17 through 22. Theradiological imaging apparatus 1C of Embodiment 6 aims at locatingunscattered γ-rays with high efficiency and generating highly accuratePET images.

[0141] The radiological imaging apparatus 1C employs the same hardwareconfiguration as the radiological imaging apparatus 1B shown in FIG. 10,except that the former uses an image pickup device 2C in place of theimage pickup device 2B and a signal processor 7B in place of the signalprocessor 7A. The image pickup device 2C has the same configuration asthe image pickup device 2B, except that the former uses an X-ray sourcecircumferential transfer unit 37C in place of the calibrated radiationsource circumferential transfer unit 37B and does not contain thecalibrated radiation source 31. The other components of the image pickupdevice 2C are the same as those of the image pickup device 2B. The X-raysource circumferential transfer unit 37C includes an X-ray source drive30C, an X-ray source device 29C that is provided with an arm 38 and anX-ray source 60, and a guide rail 28. The X-ray source drive 30C has thesame configuration as the calibrated radiation source drive 30. TheX-ray source 60 has the same configuration as the counterpart ofEmbodiment 3. The signal processor 7B has the same configuration as thesignal processor 7A, except that the former uses a coincidence counter9A in place of the coincidence counter 9 as shown in FIG. 18. Theradiological imaging apparatus 1C includes a drive controller 35 and aradiation source controller 69. Sets of three layered radiationdetectors 4 are arranged in the direction of the radius of athrough-hole 6 as shown in FIG. 19 and mounted on a radiation detectorsupport plate 5 as shown in FIG. 3.

[0142] After the administration of a PET pharmaceutical, a bed 16 onwhich a medical examinee 17 lies is moved to position the medicalexaminee 17 within the through-hole 6. At the beginning of an X-ray CTexamination, the drive controller 35 closes a motor switch. When themotor switch closes, a power source supplies an electrical current tothe motor (mounted in the X-ray source drive 30C) so that the X-raysource device 29C rotates around the medical examinee 17. As is the casewith Embodiment 3, the drive controller 35 also exercises switchingcontrol of a selector switch 62 (shown in FIG. 12), which is providedfor a signal discriminator 61. As is the case with Embodiment 3, theradiation source controller 69 exercises open/close control over anX-ray source switch connected to an X-ray tube of the X-ray source 60during an X-ray CT examination for the purpose of allowing the X-raysource 60 to emit X-rays (80 keV) for a first preselected period of time(e.g., 1 μsec) and inhibiting the X-ray source 60 from emitting X-raysfor a second preselected period of time. As a result, X-rays come fromthe circumference to fall on the medical examinee 17.

[0143] After X-rays pass through the medical examinee 17, they aredetected by a plurality of first radiation detectors 4 that exist withina predefined area facing nearly the X-ray source 60, which is on theopposite side of the through-hole 6. X-ray detection signals output fromthe first radiation detectors 4 (contained in a row of first-layerdetectors 4X) are conveyed via a wiring 13 and a selector switch 62.

[0144] Meanwhile, γ-rays emitted from the body of the medical examinee17 due to PET pharmaceutical administration are detected by secondradiation detectors 4 in detector row 4X and radiation detectors 4 insecond-layer detector row 4Y and third-layer detector row 4Z. The γ-raydetection signals output from the first radiation detectors 4 indetector row 4X are conveyed to a γ-ray discriminator 8 via the selectorswitch 62. The γ-ray detection signals output from the radiationdetectors 4 in detector rows 4Y and 4Z are conveyed to the associatedγ-ray discriminators 8 via the wiring 13.

[0145] In accordance with the input X-ray detection signals, the X-raysignal processors 66 calculate the X-ray detection signal intensity andoutput it to the computer 11. On the other hand, the γ-raydiscriminators 8 use the input γ-ray detection signals to output a pulsesignal that corresponds to the γ-ray energy attenuation in the radiationdetectors 4 that have generated the γ-ray detection signals. The outputpulse signal is conveyed to the coincidence counter 9A.

[0146] In accordance with pulse signals that are output from the γ-raydiscriminators 8 and entered within a preselected period of time (e.g.,within 10 nsec), the coincidence counter 9A identifies the two radiationdetectors 4 that output the detection signals for γ-rays unscattered inthe radiation detectors 4 (these γ-rays are hereinafter referred to asunscattered γ-rays), and outputs a PET image data signal, which containsthe information about the positions of the radiation detectors 4(initial incidence positions) and the direction of initial incidence, tothe computer 11 (the process will be detailed later with reference toFIG. 22). Further, the coincidence counter 9A counts the two pulsesignals that are generated due to the γ-ray detection signals enteredwithin the above-mentioned preselected period of time from twoidentified radiation detectors 4, and outputs the resulting count datato the computer 11. To remove the detection signals for γ-rays scatteredwithin the body of the medical examinee 17, the coincidence counter 9Achecks whether the total energy attenuation for the γ-ray detectionsignals generating the input pulse signals (total energy) is higher thana predefined energy threshold value (checks whether γ-rays are notscattered within the body of the medical examinee 17). If the totalenergy is equal to or lower than the threshold energy value, thecoincidence counter 9A removes the pulse signal count data based on suchγ-ray detection signals. In a nutshell, the initial incidence positionsare the positions of two radiation detectors 4 that were the first todetect paired γ-rays emitted from the body of the medical examinee 17.

[0147] The computer 11 performs processing steps 90, 91, 71, 72A, 73,and 74 shown in FIG. 20. In step 90, the computer 11 first inputs theγ-ray detection signal count data from the coincidence counter 9A, thepositional information about the associated detection points, and theX-ray detection signals from the X-ray signal processors 66. In step 91,the computer 11 stores these items of input information into the storagedevice 12. In step 71, the computer 11 uses the X-ray detection signalintensity to calculate the rate of X-ray attenuation in each voxel forthe body of the medical examinee 17. As is the case with Embodiment 3,the computer 11 also uses the attenuation rate in step 71 to generateX-ray CT image data about cross sections of the medical examinee 17.

[0148] In step 72A, the PET image data about cross sections of themedical examinee 17, including a diseased area (e.g., carcinomatouslesion), is generated. Processing steps 41, 42, and 44 through 47(processing steps shown in FIG. 5 for Embodiment 1), which are performedin step 72 for Embodiment 3, are performed in step 72 A. In Embodiment6, however, no transmission data imaging is conducted with thecalibrated radiation source 31. In step 44, therefore, the in-vivoattenuation correction count (reciprocal of linear attenuationcoefficient) is calculated according to the alternative in-vivoabsorption correction method described with reference to Embodiment 1and not in accordance with the γ-ray detection signals obtained duringtransmission data imaging. In step 73, the PET image data and X-ray CTimage data are synthesized to create synthesized tomogram data as is thecase with Embodiment 3. The created data appears on the display device130 (step 74).

[0149] It is possible that γ-rays detected by a certain radiationdetector 4 may scatter. Generated γ-rays change their travelingdirection when they are scattered. If either or both of paired γ-raysdetected by the coincidence counter 9A are scattered γ-rays, thestraight line joining the radiation detectors 4 that detected suchγ-rays does not pass the generation source for the detected γ-rays. Toincrease the reliability of PET images generated for PET examination, itis therefore necessary to accurately determine whether the obtainedγ-ray detection signals relate to scattered γ-rays or unscatteredγ-rays, and locate an increased number of radiation detectors 4 thatdetected unscattered paired γ-rays.

[0150] A major feature of the radiological imaging apparatus 1C is todetermine the positions and directions of paired γ-ray initial incidenceby performing the procedure indicated in FIG. 22. In Embodiment 6, twodifferent situations are considered. In one situation, the detectionsignals (γ-ray image pickup signals) are output by three or moreradiation detectors 4 according to the aforementioned coincidencecounter 9A during a preselected period of time (e.g., 10 nsec). In theother situation, such detection signals are output by two or fewerradiation detectors 4. A characteristic procedure is a processingprocedure for identifying the positions and directions of γ-ray initialincidence in cases where three or more γ-ray image pickup signals arepicked up by coincidence counting.

[0151] First of all, the processing procedure for determining thepositions and directions of γ-ray initial incidence in cases where threeor more radiation detectors 4 are picked up by coincidence counting willbe described. FIG. 21 shows the energy/scatter angle relationshipbetween unscattered γ-rays incident on a radiation detector 4 and γ-raysscattered within the radiation detector 4. In Embodiment 6, thecharacteristic shown in FIG. 21 is taken into account. When, forinstance, one of a pair of γ-rays is scattered within a radiationdetector 4 that achieved γ-ray detection (that is, if three or moreγ-rays sharing the same generation source exist, including scatteredγ-rays), the radiation detectors 4 that detected unscattered γ-rays(initial incidence positions) and the directions of initial incidenceare determined from the scatter angle of scattered γ-radiation whilemaking use of the data about an unscattered γ-ray paired with anunscattered γ-ray serving as a scattered γ-ray generation source.Scattered γ-rays are generated when unscattered γ-rays are scatteredwithin a radiation detector 4 that detected such unscattered γ-rays.

[0152] A case where a PET examination is conducted for cancer screeningwill be described as an example. Under normal conditions,fluoro-deoxyglucose (¹⁸FDG), which is a form of glucose that tends togather at cancer cells, is first administered to a medical examinee 17as a PET pharmaceutical. When administered, ¹⁸FDG emits positrons. Whenpositrons annihilate, a pair of γ-rays having a predefined energy (511keV when ¹⁸FDG is administered) are emitted. These unscattered pairedγ-rays emitted from the same source travel in virtually oppositedirections. The energy of unscattered γ-rays incident on a radiationdetector 4 remains at 511 keV unless they are scattered within the bodyof the medical examinee 17 or elsewhere before being incident on theradiation detector 4.

[0153] Suppose that, when ¹⁸FDG is uses as a PET pharmaceutical, pairedγ-rays having an energy of 511 keV are emitted in this manner from thebody of the medical examinee 17, one of the unscattered paired γ-rayssuffers 100 keV attenuation in radiation detector 4 g shown in FIG. 19,scattered γ-rays generated in radiation detector 4 g due to such anunscattered γ-ray suffer 100 keV attenuation in radiation detector 4 h,and the remaining unscattered paired γ-ray suffers total attenuation inradiation detector 4 f (the positions of radiation detectors 4 f, 4 g,and 4 h are regarded as O, A, and B, respectively). Radiation detectors4 f, 4 g, 4 h in which γ-rays are attenuated output γ-ray detectionsignals. In this instance, the paired γ-rays are emitted in oppositedirections due to a PET pharmaceutical gathered in a diseased area.Therefore, the direction of paired γ-ray travel (more specifically thetraveling direction of an opposite unscattered γ-ray) is a combinationof the vector OA and vector AO or a combination of the vector OB andvector BO. If the direction of an unscattered γ-ray is the vector OA(the initial incidence position is position A), the energy of ascattered γ-ray is 411 keV. If the direction of an unscattered γ-ray isthe vector OB (the initial incidence position is position B), the energyof a scattered γ-ray is 100 keV. In other words, when the initialincidence position of an unscattered γ-ray is position A, the directionof a scattered γ-ray is the vector AB. If the initial incidence positionis position B, the direction of a scattered γ-ray is the vector BA. Theunscattered γ-ray energy is equal to the sum of both energyattenuations, that is 411+100=511 keV, no matter whether initialincidence occurs at position A or position B. The above energyattenuation values depend on the pulse height of a pulse signal outputfrom a γ-ray discriminator 8. It can therefore be said that the energyattenuation values are detected by the associated γ-ray discriminators8.

[0154] In Embodiment 6, the coincidence counter 9A performs processingsteps 92 through 97 shown in FIG. 22 to determine the positions anddirections of γ-ray initial incidence, using the energy attenuationvalues detected in relation to the above γ-rays and the positionalinformation about radiation detectors 4 on which such γ-rays areincident. The positional information about radiation detectors 4 thatoutput γ-ray detection signals is converted to a pulse signal by theγ-ray discriminators 8 provided for the radiation detectors 4 andconveyed to the coincidence counter 9A. In step 92, the coincidencecounter 9A first determines a radiation detector candidate thatgenerated a scattered γ-ray. When three or more pulse signals areentered during the aforementioned preselected period of time, thecoincidence counter 9A checks three or more radiation detectors 4 thatoutput γ-ray detection signals, which have caused the generation of thepulse signals, and then determines a radiation detector that hasgenerated a scattered γ-ray. This determination is made in accordancewith the distance between the associated radiation detectors 4. In otherwords, radiation detectors 4 positioned at a spacing interval shorterthan preselected are the radiation detectors 4 that have generatedscattered γ-rays. Since the scattered γ-ray transmission distance isshort, suppose that the distance setting is 5 cm. In the example shownin FIG. 19, in which γ-ray detection signals are output within apreselected period of time by radiation detectors 4 f, 4 g, and 4 h, ifthe distance between the radiation detectors 4 g and 4 h is 5 cm orshorter, scattered γ-rays are generated by either the radiation detector4 g or the radiation detector 4 h. The radiation detectors 4 g and 4 hare determined as radiation detector candidates that have generatedscattered γ-rays. Therefore, the coincidence counter 9A recognizes thatan unscattered γ-ray is detected by the remaining radiation detector 4f.

[0155] In step 93, which is performed after the radiation detectors 4 gand 4 h are designated as radiation detector candidates that havegenerated scattered γ-rays, the scatter angle (angle formed by thevectors OA and AB) prevailing when the radiation detector 4 g, that is,position A is the initial incidence position is calculated. In step 94,the scatter angle (angle formed by the vectors OB and BA) prevailingwhen radiation detector 4 h, that is, position B is the initialincidence position is calculated. In this instance, the angle θ, whichis formed by the vectors OA and AB, can be calculated as indicatedbelow:

θ=⁻¹({right arrow over (OA)}·{right arrow over (AB)})/(|{right arrowover (OA)}|·|{right arrow over (AB)}|)  (3)

[0156] The next step, which is step 95, is performed to calculate theunscattered γ-ray incidence energy and scattered γ-ray energy. Morespecifically, the coincidence counter 9 calculates the γ-ray energyattenuations in the radiation detectors 4 g and 4 h in accordance withthe pulse heights of pulse signals generated due to individual γ-raydetection signal outputs from the radiation detectors. The energyattenuation of a γ-ray incident on radiation detector 4 f in the initialincidence position is calculated to be 511 keV according to the pulseheight of the associated pulse signal. Further, the energy attenuationof a γ-ray incident on the radiation detector 4 g is calculated to be100 keV according to the pulse height of the associated pulse signal. Inlike manner, the energy attenuation of a γ-ray incident on radiationdetector 4 h is calculated to be 411 keV. A scattered γ-ray is generatedby the radiation detector 4 g or 4 h. An unscattered γ-ray is detectedby the radiation detector 4 g or 4 h, whichever did not generate thescattered γ-ray. The energy of unscattered γ-ray incidence is the sum ofenergy attenuations in the radiation detectors 4 g and 4 h andcalculated to be 511 keV.

[0157] Next, the scattered γ-ray energy is calculated on the presumptionthat scattered γ-rays are generated in both of the radiation detectors 4g and 4 h. When a scattered γ-ray is generated in the radiation detector4 g, the energy of that scattered γ-ray is 411 keV (=511 keV−100 keV).In this instance, the scattered γ-ray eventually attenuates in theradiation detector 4 h. When a scattered γ-ray is generated in theradiation detector 4 h, the energy of that scattered γ-ray is 100 keV(=511 keV−411 keV). In this instance, the scattered γ-ray eventuallyattenuates in the radiation detector 4 f. If, for instance, thecalculated energy of an unscattered γ-ray is considerably lower than 511keV (e.g., below 350 keV), such a ray is excluded because it isconceivable that it was previously scattered within the body of themedical examinee 17.

[0158] When the initial incidence position is either position A orposition B, steps 93 through 95 are performed to calculate the incidentγ-ray energy, scattered γ-ray energy, and scatter angle. Step 96 isperformed to determine what attenuation sequence (scattering sequence)is appropriate by checking whether the calculated relationship amongincidence γ-ray energy, scattered γ-ray energy, and scatter angle agreeswith the relationship indicated in FIG. 21. If, for instance, a scatterangle comparison is to be made, the incident γ-ray energy and scatteredγ-ray energy indicated in FIG. 21 are used to calculate an ideal scatterangle and determine the deviation of the actual scatter angle from theideal one. An appropriate threshold value (e.g., for tolerating adeviation of up to 10%) is set for the relationship between calculationresults and FIG. 21. When the threshold value is not exceeded by thedeviation, it is concluded that an expected phenomenon can occur (theattenuation sequence is proper). When it is concluded that one of two ormore phenomena (two phenomena in the currently used example) can occur,such a phenomenon is selected. (Cases where it is concluded that two ormore phenomena can occur will be explained later.) As a result, theinitial incidence positions of paired γ-rays are determined. Finally,step 91 is performed to output a PET image data signal, which containsthe information about the initial incidence positions of unscatteredpaired γ-rays and the straight line (direction of initial incidence)joining these positions, to the computer 11. The procedure is nowcompleted.

[0159] As described earlier, the computer 11 stores a large number ofPET image data, which is entered in the above manner, into the storagedevice 12, reconstructs them to formulate a PET image, and displays iton the display device 130.

[0160] If it concluded after the above procedure is performed in caseswhere three or more detection signals are counted that the initialincidence position can be either position A or position B, analternative is to remove the associated γ-ray detection signals orselect an attenuation sequence so as to minimize the deviation from therelationship indicated in FIG. 21. In the above-described situation, oneof paired γ-rays is attenuated (subjected to two attenuations) by tworadiation detectors 4 (this example is simple because either of twodifferent attenuation sequences is to be selected). However, it is alsopossible to determine a proper attenuation sequence by performing theabove-described procedure for all possible patterns, including the onein which generated paired γ-rays are both scattered a number of timeswithin the radiation detectors 4. Another alternative is to define thepositional relationship between the radiation detectors 4 that cannotphysically be subjected to coincidence counting due to the layout of theradiation detectors 4, and remove a pulse signal that corresponds to thedefined positional relationship.

[0161] Meanwhile, if two γ-ray detection signals are simultaneouslycounted, the coincidence counter 9A concludes that they relate to pairedγ-rays, and outputs the positional information about radiation detectors4 that detected such rays and the PET image data including theinformation about the direction of a straight line joining suchradiation detectors to the computer 11. Alternatively, a differentconfiguration may be used to achieve γ-ray detection signal removal whenonly one γ-ray detection signal is counted. In another alternativeconfiguration, a data signal containing the information about thedirection of a straight line joining the radiation detector 4 thatachieved γ-ray detection to an opposing radiation detector, which ispositioned 180° apart, may be output to the computer 11 in a mannersimilar to a conventional one.

[0162] Embodiment 6 provides the following advantages in addition toadvantages (1) through (13) of Embodiment 3.

[0163] (14) PET Image Accuracy Enhancement

[0164] Embodiment 6 makes it possible to effectively determine thepositions and directions of paired γ-ray initial incidence by performinga predetermined procedure indicated in FIG. 22. As a result, highlyreliable data can be output to the computer 11 to enhance the PET imageaccuracy. Although ¹⁸FDG is used as a PET pharmaceutical for theexplanation of Embodiment 6, the aforementioned processing procedure fordetermining the positions and directions of γ-ray initial incidence isalso applicable to cases where the employed PET pharmaceutical containsthe other radionuclides.

[0165] (15) PET Image Accuracy Enhancement

[0166] In Embodiment 6, the positional information about radiationdetectors that have detected a γ-ray and the positional informationabout radiation detectors that have detected the other γ-ray are used toset up possible attenuation sequences of the former γ-ray. These γ-rayattenuation sequences are examined to select an appropriate one, whichexhibits a proper relationship between the γ-ray scatter angle andenergy detection value. In this manner, the γ-ray attenuation sequenceis determined. As a result, the position of initial γ-ray incidence on aradiation detector is determined. It can therefore be concluded that aγ-ray generation source (diseased area) exists on a straight line(direction of initial incidence) joining the determined radiationdetector to a radiation detector that detected the other γ-ray. Unlikeprobabilistic determination of initial γ-ray incidence position,unscattered γ-rays can be determined with high efficiency to generate ahighly accurate PET image.

[0167] In particular, when the γ-ray detection signals of three or moreradiation detectors are simultaneously counted, Embodiment 6 uses thepositional information about the three or more γ-ray detection signalsto set up possible γ-ray attenuation sequences, and selects anappropriate sequence that exhibits a proper relationship to energydetection values from the three or more radiation detectors. The initialγ-ray incidence position determined in this manner and the above energydetection value data can then be used to determine the direction ofinitial γ-ray incidence. Unlike probabilistic determination of initialγ-ray incidence position, unscattered γ-rays can be determined with highefficiency to generate a highly accurate PET image.

[0168] (16) Incorporation of X-ray CT Examination and PET ExaminationFunctions

[0169] In the past, an image pickup device for detecting transmittedX-rays was generally installed independently of an image pickup devicefor detecting γ-rays. However, Embodiment 6 uses the same radiationdetectors 4 to detect X-rays and γ-rays. Therefore, the aforementionedimage pickup device 2 can provide both X-ray CT examination and PETexamination although it has a simple, compact structure.

[0170] Embodiment 7

[0171] A radiological imaging apparatus of another embodiment(Embodiment 7) of the present invention will be described with referenceto FIGS. 23 and 24. The radiological imaging apparatus 1D of Embodiment7 is a two-dimensional measurement type radiological imaging apparatus.Embodiment 7 determines the γ-ray attenuation sequence from the energiesand scatter angles of unscattered γ-rays and scattered γ-rays, andprovides a detection efficiency improvement over single γ-ray detectionby a two-dimensional measurement type radiological imaging apparatus. Intwo-dimensional measurements, the apparatus uses a collimator to removeγ-rays that are initially incident at a certain angle on the directionof the radiation detector body axis (equivalent to the axial directionof the aforementioned through-hole 6), and detects only γ-rays that areinitially incident perpendicularly to the direction of the body axis.Two-dimensional measurements, in which γ-rays incident at a certainangle are removed, generally decrease the γ-ray pair count per unittime, but offer an advantage that the influence of scattered γ-rays canbe minimized.

[0172] The radiological imaging apparatus 1D has the same configurationas the radiological imaging apparatus 1C, except that an image pickupdevice 2D of the radiological imaging apparatus 1 d includes acollimator 98. The collimator 98 is mounted on a detector support plate5 and positioned in front of (inner circular side of) radiationdetectors 4 in the innermost detector row 4X (see FIG. 18).

[0173] For PET examination, X-ray CT examination, and synthesizedtomogram data creation, the radiological imaging apparatus 1D uses theprocedures as the radiological imaging apparatus 1C. However, theprocess performed by the radiological imaging apparatus 1D fordetermining the positions and directions of initial γ-ray incidence isdifferent from that which is performed by the radiological imagingapparatus 1C. Embodiment 7 determines the γ-ray attenuation sequencewhen γ-rays incident radiation detectors 4 after passing through thecollimator 98 are scattered three or more times within a radiationdetector 4. For the sake of brevity, Embodiment 7 is described here onthe presumption that γ-rays incident on radiation detectors 4 totallyattenuate in radiation detectors 4 i, 4 j, and 4 k in an arbitrary orderas indicated in FIG. 24. However, the positions of the radiationdetectors 4 i, 4 j, and 4 k are regarded as positions A, B, and C,respectively, and the energies attenuated at positions A, B, and C areregarded as E_(A), E_(B), and E_(C), respectively. For the sake ofconvenience, the drawings depicting Embodiment 7 are prepared on thepresumption that positions A, B, and C are in the same plane. However,the following description of Embodiment 7 is also applicable to caseswhere positions A, B, and C are not in the same plane as indicated inFIG. 24.

[0174] From FIG. 24, six different attenuation sequences areconceivable: {circle over (1)} B→A→C, {circle over (2)} C→A→B, {circleover (3)} A→B→C, {circle over (4)} C→B→A, {circle over (5)} A→C→B, and{circle over (6)} B→C→A. Further, since total attenuation occurs atpositions A, B, and C, the γ-ray energy prevailing at the time ofinitial incidence (total energy E) is E_(A)+E_(B)+E_(C). Therefore, theenergy of scattered γ-ray incident on the second attenuation position isdetermined by subtracting the first energy attenuation (E_(A), E_(B), orE_(C)) from the total energy E. The energy of the γ-ray scattered at thesecond position is the energy of scattered γ-ray incident on the thirdattenuation position (E_(A), E_(B), or E_(C)).

[0175] Therefore, the scattered γ-ray scatter angel at the secondattenuation position and the incoming energy and outgoing energy asviewed from the second attenuation position are calculated with respectthe above-mentioned six different attenuation sequences. The obtainedresults are then compared with the relationship indicated in FIG. 7 tocheck the six different attenuation sequences to determine a properattenuation sequence that can occur in reality.

[0176] Next, the scatter angle at the initial incidence position isdetermined in accordance with the attenuation sequence determined inthis manner. If, for instance, the attenuation sequence {circle over(3)} (A→B→C) is found to be proper as a result of the check of the sixdifferent attenuation sequences, the incidence energy prevailing at thefirst attenuation position A is E_(A)+E_(B)+E_(C), and the outgoingenergy is E_(B)+E_(C). When these energy values compared with therelationship indicated in FIG. 21, the scatter angle prevailing atposition A is identified. When considering the fact that only γ-raysperpendicular to the direction of the body axis are incident on theradiation detectors 4 in a two-dimensional measurement PET examination,it is concluded that the direction of initial γ-ray incidence isindicated by arrow 99 a or 99 b in FIG. 24. As is obvious from FIG. 24,the existence range of a γ-ray generation source does not physicallyallow the arrow 99 b to represent the direction of initial γ-rayincidence. Therefore, it is uniquely concluded that the γ-rays areinitially incident on the radiation detector 4 g (initial incidenceposition) and that the initial incidence direction is as indicated bythe arrow 99 a.

[0177] If no more than two γ-ray detection signals were simultaneouslycounted (that is, if the attenuation in radiation detectors did notoccur more than two times), the associated data is removed and not used.An alternative is to determine the first attenuation position on thepresumption that uniform incidence occurs from the range of radiationsource existence.

[0178] Embodiment 7 uses a coincidence counter 9A to determine thepositions and directions of initial γ-ray incidence. As describedearlier, a γ-ray discriminator 8 converts a γ-ray detection signalhaving an energy value not less than a predefined threshold value to apulse signal, and outputs it to the coincidence counter 9A. In thisinstance, not only the pulse signal but also the positional informationabout a radiation detector 4 whose γ-ray detection signal is detected isoutput to the coincidence counter 9A. The coincidence counter 9Adetermines the position and direction of γ-ray initial incidence inaccordance with the pulse signal input from the γ-ray discriminator 8,and outputs a PET image data signal, which contains the informationabout the position and direction of γ-ray initial incidence, to thecomputer 11. When three or more pulse signals are simultaneouslycounted, Embodiment 7 causes the coincidence counter 9A to perform theabove procedure for determining the position and direction of initialincidence. In the other situations, however, the coincidence counter 9Aperforms the following procedures depending on the encounteredsituation.

[0179]FIG. 25 shows typical input and output signals of the coincidencecounter 9A. The numbers parenthesized in FIG. 25 indicate the number ofsignal inputs or outputs. If, for instance, the position and directionof γ-ray initial incidence are determined from an input pulse signal asin a case i, v, or vi shown in FIG. 25, the coincidence counter 9Aoutputs a PET image data signal, which contains the information aboutthe determined position and direction of initial incidence, to thecomputer 11. If, for instance, there is no pulse signal input (case i),one totally-attenuated γ-ray pulse signal is entered (case ii), or threepulse signals, which cannot possibly be generated from the same sourcedue to the layout of radiation detectors 4, are counted (case vii), thecoincidence counter 9A remove the pulse signal(s) and does not output aPET image data signal. If two totally-attenuated γ-ray pulse signals arecounted (case iv), the coincidence counter 9A outputs a data signal,which contains the positional information about radiation detectors 4whose signals are detected and a straight line joining these radiationdetectors, to the computer 11.

[0180] As is the case with Embodiment 6, the computer 11 stores theinput PET image data signal in the storage device 12. The count data forthe aforementioned γ-ray detection signals are also stored into thestorage device 12 by the computer 11. If three or more signals arecounted (case vii) and the direction of incidence is known, the data forthat direction may be output. The data obtained in this manner isreconstructed by the computer 11 and displayed on the display device130.

[0181] Embodiment 7 provides some advantages in addition to theadvantages of Embodiment 6. When either or both of paired γ-rays arescattered, Embodiment 6 determines the attenuation sequence of atargeted unscattered γ-ray in accordance with the detection signal forthe remaining unscattered γ-ray. However, Embodiment 7 can consider thescatter status of a paired γ-ray and determine the initial incidenceposition and direction (that is, initial incidence direction) of anunscattered γ-ray during a two-dimensional measurement PET examinationeven when the remaining paired γ-ray is totally attenuated (absorbed)within the body of a medical examinee. In this manner, Embodiment 7 cancollect the data about unscattered γ-rays with high efficiency andincrease the PET image accuracy. As a result, the count per unit time oftwo-dimensional measurement PET examination increases, making itpossible to reduce the examination time. It can also be expected thatthis advantage will reduce the load on a medical examinee 17 andincrease the throughput of the number of medical examinees. If theincidence of each of paired γ-rays is verified, the scattered γ-rayattenuation sequence determination procedure performed for Embodiment 6is also applicable to a two-dimensional measurement type PET examinationapparatus described according to Embodiment 7.

[0182] In an X-ray CT image creation example that is described accordingto Embodiments 6 and 7, the arm 38 is sequentially extended andcontracted to create tomograms of various cross sections of a medicalexaminee 17. However, when the X-ray source 60 is rotated simultaneouslywith the extension/contraction of the arm 38, Embodiments 6 and 7 areapplicable to an X-ray helical scan as well. Further, an alternativeconfiguration may be employed so that the bed 16 moves in the axialdirection of the through-hole 6 instead of the extension/contraction ofthe arm 38.

[0183] The above PET/X-ray examination procedure may be performed inrelation to the entire body of a medical examinee 17 or in relation tothe neighborhood of the medical examinee's diseased area roughly locatedbeforehand by another examination. In some situations, the examinationmay be conducted without administering a PET pharmaceutical to themedical examinee 17 in advance but administering the PET pharmaceuticalto the medical examinee 17 laid on the bed 16, or conducted whileadministering the PET pharmaceutical to the medical examinee 17. Analternative configuration, which is not specifically described withreference to the above first and second embodiments, may be employed sothat the radiological imaging apparatus 1C, 1D is provided with aseparate calibrated radiation source to perform transmission imaging.These alternative embodiments also provide the same advantages.

[0184] In Embodiments 6 and 7, a predetermined procedure is followed sothat the coincidence counter 9A determines the γ-ray attenuationsequence, initial incidence position, and initial incidence direction.Alternatively, however, a separate circuit performing this process maybe furnished to complete the process at high speed. Another alternativescheme may be used so that a coincidence counter circuit merely selectssimultaneous events, allowing the software to carry out the subsequentprocessing steps. That is, when, for instance, three signals areentered, the coincidence counter 9A sends information, which containsthe data for indicating that the three signals are coincident, to thecomputer 11, and the computer 11 performs a predetermined procedure todetermine the attenuation sequence, initial incidence position, andinitial incidence direction. Even if each radiation detector is providedwith a storage area, the γ-ray incidence time and the γ-ray energyattenuation in a radiation detector are written in such an area, and thecomputer 11 reads the written data to check for coincidence, thecomputer 11 can determine the attenuation sequence, initial incidenceposition, and initial incidence direction by performing a predeterminedprocedure. In this instance, also the same advantages are obtained.

[0185] In Embodiments 6 and 7, the incidence scatter angle can bedetermined even if one of paired γ-ray is scattered in a radiationdetector and the remaining paired γ-ray is not detected by a radiationdetector. If it is known that the γ-rays are emitted from a certain area(e.g., from within a plane), the above property can be used to determinewhich of the areas into which the γ-rays can enter is the source ofgeneration. These data can be effectively used to raise the detectionefficiency of radiation detectors and reduce the load on patients.

[0186] In Embodiments 6 and 7, the radiation detectors 4 in multiplelayers are linearly arranged in the radial direction with the innermostones regarded as the base points as shown in FIGS. 19 and 24.Alternatively, however, the radiation detectors 4 may be zigzagged inthe direction of the radius. Although the above descriptions deal with aPET examination in which emitted paired γ-rays are to be detected, it isalso possible that α- and γ-rays or β- and γ-rays may be paired whenemitted. In these instances, the above-described attenuation sequencedetermination procedure works because γ-rays may scatter multiple timesalthough the α- and β-rays have a low penetrating power. Although theabove descriptions deal with cases where the coincidence counter 9Adetermines the positions of γ-ray initial incidence, it is alternativelypossible that the position, energy detection value, and detection timedata about the radiation detectors 4 may be output to the computer 11 toallow the computer 11 to perform the above-described processingprocedure. In all the above cases, also the same advantages areobtained.

What is claimed is:
 1. A radiological imaging apparatus comprising: animage pickup device that is provided with a plurality of radiationdetectors for detecting the radiation from a subject, wherein theradiation passing through one of said radiation detectors is to bedetected by another one of said radiation detectors; and signalprocessors that are each connected to corresponding one of saidradiation detectors to process a radiation detection signal detected bysaid corresponding one on said radiation detectors.
 2. The radiologicalimaging apparatus according to claim 1, wherein said image pickup deviceis provided with an X-ray source for irradiating said subject withX-rays and a flat panel detector equipped with said plurality ofradiation detectors.
 3. The radiological imaging apparatus according toclaim 1, further comprising a bed on which said subject is to be placed.4. The radiological imaging apparatus according to claim 1, wherein saidone of said radiation detectors and said another one of said radiationdetectors are each mounted on a radiation detector support member. 5.The radiological imaging apparatus according to claim 1, wherein saidimage pickup device includes an X-ray source that moves around saidsubject to irradiate said subject with X-rays.
 6. The radiologicalimaging apparatus according to claim 1, wherein said radiation detectorsdetect γ-rays that are emitted from said subject due to aradiopharmaceutical administered to said subject.
 7. The radiologicalimaging apparatus according to claim 6, wherein said image pickup deviceincludes an X-ray source that moves around said subject to irradiatesaid subject with X-rays, and said radiation detectors also detectX-rays that pass through said subject after being emitted from saidX-ray source.
 8. The radiological imaging apparatus according to claim3, wherein said one of said radiation detectors and said another one ofsaid radiation detectors are linearly arranged.
 9. A radiologicalimaging apparatus comprising: a bed on which a medical examinee is to belaid; an image pickup device that includes a plurality of radiationdetectors for detecting the radiation from said medical examinee, saidradiation detectors being disposed in said image pickup device, arrangedaround the circumference of a through-hole into which said bed is to beinserted, and mounted at different positions in the radial direction ofsaid through-hole; and signal processors that are each connected tocorresponding one of said radiation detectors to process a radiationdetection signal detected by said corresponding one of said radiationdetectors.
 10. The radiological imaging apparatus according to claim 9,wherein said radiation detectors are mounted on radiation detectorsupport members that are positioned around the circumference of saidthrough-hole.
 11. The radiological imaging apparatus according to claim9, wherein said plurality of radiation detectors mounted at differentpositions in the radial direction of said through-hole are arrangedlinearly in said radial direction.
 12. The radiological imagingapparatus according to claim 9, further comprising a tomogram generatorfor receiving information output from said signal processors andcreating tomogram data about said subject in accordance with thereceived output information.
 13. The radiological imaging apparatusaccording to claim 12, wherein said information output from said signalprocessors is obtained when γ-ray detection signals, which are saidradiation detection signals, are processed by said signal processors.14. The radiological imaging apparatus according to claim 9, whereinsaid image pickup device includes a γ-ray source that moves around saidsubject to irradiate said subject with γ-rays, and said radiationdetectors detect a first γ-ray that passes through said subject afterbeing emitted from said γ-ray source as well as a second γ-ray that isemitted from said subject due to a radiopharmaceutical administered tosaid subject.
 15. The radiological imaging apparatus according to claim14, further comprising a tomogram data generator; wherein said signalprocessors output first information upon receipt of first γ-raydetection signals that are output from said radiation detectors upondetection of said first γ-ray and output second information upon receiptof second γ-ray detection signals that are output from said radiationdetectors upon detection of said second γ-ray; and wherein said tomogramdata generator corrects said second information in accordance with saidfirst information and creates tomogram data about said subject from saidcorrected second information.
 16. The radiological imaging apparatusaccording to claim 9, wherein said image pickup device includes an X-raysource that moves around said subject to irradiate said subject withX-rays, wherein said plurality of radiation detectors form multiplelayers of radiation detectors in said radial direction, and wherein saidradiation detectors included at least in the first layer from saidthrough-hole output a first detection signal, which is the detectionsignal for one of said rays, that is, said X-ray passing through saidsubject as well as a second detection signal, which is the detectionsignal for another one of said rays, that is, a γ-ray emitted from saidsubject.
 17. The radiological imaging apparatus according to claim 16,further comprising a tomogram data generator for creating first tomogramdata about said subject in accordance with first information that isoutput from said signal processors upon input of said first detectionsignal, creating second tomogram data about said subject in accordancewith second information that is output from said signal processors uponinput of said second detection signal, and creating synthesized tomogramdata by synthesizing said first tomogram data and said second tomogramdata.
 18. The radiological imaging apparatus according to claim 16,comprising a tomogram data generator; wherein said signal processorsconnected to said radiation detectors in some radiation detector layers,positioned on said through-hole side, of said multiple radiationdetector layers include X-ray signal processors for outputting firstinformation about X-rays in accordance with said first detection signaland γ-ray signal processors for outputting second information aboutγ-rays in accordance with said second detection signal, and said signalprocessors connected to said radiation detectors in the remainingradiation detector layers of the multiple radiation detector layers donot include said X-ray signal processors but include said γ-ray signalprocessors; and wherein said tomogram data generator creates firsttomogram data about said subject in accordance with said firstinformation that is output from said X-ray signal processors, createssecond tomogram data about said subject in accordance with said secondinformation that is output from said γ-ray signal processors, andcreates synthesized tomogram data by synthesizing said first tomogramdata and said second tomogram data.
 19. The radiological imagingapparatus according to claim 9, further comprising a radiation detectordeterioration check device; wherein said plurality of radiationdetectors are arranged in said direction of the radius to form multipleradiation detector layers, wherein said signal processors connected tosaid radiation detectors output first information upon input of a γ-raydetection signal, which is said radiation detection signal; and whereinsaid radiation detector deterioration check device determines themeasured detection efficiency proportion of radiation detectors in saidmultiple layers in accordance with said first information about theradiation detectors and uses the resultant measured value proportion andtheoretical detection efficiency proportion of the radiation detectorsto determine whether the radiation detectors are deteriorated.
 20. Theradiological imaging apparatus according to claim 9, wherein saidradiation detectors are semiconductor radiation detectors.
 21. Theradiological imaging apparatus according to claim 1, further comprisinga counter; wherein said counter uses the γ-ray detection signals outputfrom three or more of said plurality of radiation detectors within apreselected period of time and the positional information about saidthree or more radiation detectors that have output the γ-ray detectionsignals, so as to determine which of said three or more radiationdetectors have detected unscattered γ-rays in said radiation detectors.22. A radiological imaging apparatus comprising: a plurality ofradiation detectors for detecting γ-rays; and a counter that uses theγ-ray detection signals output from three or more of said plurality ofradiation detectors within a preselected period of time and thepositional information about said three or more radiation detectors thathave output the γ-ray detection signals, so as to determine which ofsaid three or more radiation detectors have detected unscattered γ-raysin said radiation detectors.
 23. A radiological imaging apparatuscomprising: a plurality of radiation detectors for detecting γ-raysemitted from a subject to which a radiopharmaceutical is administered;and a counter that, when γ-ray detection signals are output from threeor more of said plurality of radiation detectors within a preselectedperiod of time, uses the positional information about at least two ofsaid radiation detectors, the energy detection values of at least two ofsaid radiation detectors, and the positional information about radiationdetectors that have detected one of a pair of said γ-rays, so as todetermine the attenuation sequence, initial incidence position, andinitial incidence direction of the remaining one of said pair of γ-rays.24. The radiological imaging apparatus according to claim 23, whereinsaid attenuation sequence, said initial incidence position, and saidinitial incidence direction are determined by checking two or moredifferent attenuation sequences of one of said γ-rays, which can beestimated from the positional information about radiation detectors thathave detected one of said γ-rays and radiation detectors that havedetected the other one of said γ-rays, and selecting a sequenceexhibiting a proper relationship between the scatter angle and energydetection value of one of said γ-rays.
 25. A radiological imagingapparatus comprising: a plurality of radiation detectors for detectingγ-rays; collimators positioned in front of said plurality of radiationdetectors to permit γ-ray passage; and a counter that, when detectionsignals are output from three or more of said plurality of radiationdetectors within a preselected period of time, uses the positionalinformation about three or more of said radiation detectors and theenergy detection values of three or more of said radiation detectors inorder to determine the attenuation sequence, initial incidence position,and initial incidence direction of said γ-rays.
 26. The radiologicalimaging apparatus according to claim 25, wherein said attenuationsequence, said initial incidence position and said initial incidencedirection are determined by checking two or more different attenuationsequences of said γ-rays, which can be estimated from said positionalinformation, and selecting a sequence exhibiting a proper relationshipto said energy detection value.
 27. The radiological imaging apparatusaccording to claim 26, wherein the initial incidence position of saidγ-rays is determined in accordance with said selected proper attenuationsequence, and then the determined initial incidence position and saidenergy detection value are used to determine the initial incidencedirection of said γ-rays.
 28. The radiological imaging apparatusaccording to claim 22, wherein said plurality of radiation detectors arearranged in the form of a ring while a number of said plurality ofradiation detectors are arrayed in the axial direction and in multiplelayers in the radial direction.
 29. The radiological imaging apparatusaccording to claim 28, further comprising a γ-ray discriminator thatoutputs a pulse signal upon receipt of a γ-ray detection signal inputfrom said radiation detector.
 30. The radiological imaging apparatusaccording to claim 29, wherein said counter outputs the positionalinformation about said determined radiation detector and the countinformation about said pulse signal.
 31. The radiological imagingapparatus according to claim 30, further comprising a tomogram generatorfor generating tomogram data in accordance with the positionalinformation about said radiation detectors and said count informationand a display device for displaying said tomogram data.
 32. Theradiological imaging apparatus according to claim 22, further comprisinga tomogram generator for generating tomogram data in accordance with thepositional information about said determined radiation detector and theγ-ray detection signal output from said determined radiation detector.33. The radiological imaging apparatus according to claim 28, furthercomprising an X-ray source for emitting X-rays.
 34. The radiologicalimaging apparatus according to claim 33, further comprising a signaldiscriminator for discriminating the detection signals for γ-rays andX-rays that are detected by a plurality of shared radiation detectors,which are among said multiple-layered radiation detectors, mounted in atleast the innermost area, and used for detecting γ-rays and X-rays.